Implantable, retrievable, thrombus minimizing sensors

ABSTRACT

A sensor is disclosed, for implantation within a blood vessel to monitor a substance in or property of blood. In one embodiment, the sensor is retrievable. In another embodiment, the sensor has a layer that minimizes the formation of thrombus. A signal representative of the substance in or property of blood is transmitted to an external receiver. In another embodiment, the sensor detects nitric oxide or a nitric oxide metabolite and can be implanted for a period of hours, days, weeks or years. Methods are also disclosed.

BACKGROUND OF THE INVENTION Field of the Invention

This is a continuation-in-part of U.S. patent application Ser. No.10/217,202, filed Aug. 9, 2002 now U.S. Pat. No. 7,006,858, which is acontinuation-in-part of U.S. patent application Ser. No. 10/041,036,filed Nov. 8, 2001 now U.S. Pat. No. 7,033,322, which is acontinuation-in-part of U.S. patent application Ser. No. 09/571,702,filed May 15, 2000, now U.S. Pat. No. 6,442,413 issued on Aug. 27, 2002.

DESCRIPTION OF THE RELATED ART

The present invention generally relates to the use of sensors to monitorthe concentration of a chemical species in bodily fluids. Morespecifically, the present invention relates to the use of sensors tomonitor glucose levels, and/or other parameters in a fluid, includingpressure or flow rate within a lumen of an endoluminal implant such as astent or other type of endovascular conduit.

Diabetes mellitus is a serious medical condition affecting approximately10.5 million Americans, in which the patient is not able to maintainblood glucose levels within the normal range (normoglycemia).Approximately 10% of these patients have insulin-dependent diabetesmellitus (Type I diabetes, IDDM), and the remaining 90% havenon-insulin-dependent diabetes mellitus (Type II diabetes, NIDDM). Thelong-term consequences of diabetes include increased risk of heartdisease, blindness, end-stage renal disease, and non-healing ulcers inthe extremities. The economic impact of diabetes to society has beenestimated by the American Diabetes Association at approximately $45.2billion annually (Jonsson, B., The Economic Impact of Diabetes, DiabetesCare 21(Suppl 3): C7–C10, (1998)).

A major long-term clinical study, the Diabetes Control and ComplicationsTrial, involving 1,441 patients with insulin-dependent diabetes mellitus(Type I diabetes) over a 10-year period from 1984–1993, demonstratedthat by intensive therapy (frequent administration of either short- orlong-acting insulin), these long-term consequences (retinopathy,nephropathy, and neuropathy) could be reduced (“The Effect of IntensiveTreatment of Diabetes on the Development and Progression of Long-TermComplications in Insulin-Dependent Diabetes Mellitus,” The DiabetesControl and Complications Trial Research Group, New Eng. J. Med., 329:977–86 (1993)). Unfortunately, a major difficulty encountered during thetrial was that intensive treatment also resulted in a higher incidenceof low blood glucose levels (hypoglycemia), which was severe enough toresult in coma or death, as compared to patients under conventionalmedical management.

Currently, diabetics must monitor their condition by repeatedly prickingtheir fingers in order to obtain blood samples for evaluation. The majordrawback to self-monitoring of glucose is that it is discontinuous andtherefore the number of glucose measurements performed is dependent onthe motivation of the patient.

Existing analytical techniques and devices for in vitro glucosemeasurements have a high level of accuracy (the error can be <1%). Manyof these routine methods are accepted as standards of comparison withnew devices. Management of diabetes currently relies on these methods tocontrol the disease and minimize complications.

There are two main disadvantages to these existing options. First,sampling even a minimal amount of blood multiple times per day isassociated with risks of infection, nerve and tissue damage, anddiscomfort to the patients. Second, in the case of dynamic changes inglucose concentration, very frequent or even continuous measurements ofblood glucose levels are required (Wilkins, E., et al., “GlucoseMonitoring: State of the Art and Future Possibilities”, Med. Eng. Phys.18(4):273–88, (19916).

There are two main approaches to the development of a continuous bloodglucose monitor. The first category is non-invasive sensors, whichobtain information from physico-chemical characteristics of glucose(spectral, optical, thermal, electromagnetic, or other). The secondcategory is invasive sensors. In this group, there is intimatemechanical contact of the sensor with biological tissues or fluids,because the device is placed within the body. (Wilkins, 1996).

Non-invasive sensor technology has focused on the absorption of thenear-infrared (NIR) spectra by the analyte, in this case, glucose (SeeU.S. Pat. No. 5,945,676 to Khalil, et al., and U.S. Pat. No. 5,433,197to Stark). Absorptions which occur in the NIR region are most oftenassociated with overtone and combination bands of the fundamentalvibrations of —OH, —NH, and —CH functional groups. As a result, mostbiochemical species will exhibit some absorption in the region ofinterest. Glucose measurements are usually performed in the spectraregion from 4250 to 660 cm⁻¹. These highly overlapping, weakly absorbingbands were initially thought to be too complex for interpretation andtoo weak for practical application. Improvements in instrumentation andadvances in multivariate chemometric data analysis techniques may allowmeaningful results to be obtained from these complex spectra.

However, to date these devices are not particularly accurate even in thenormal physiological range. A subject-dependent concentration bias hasbeen reported. The temperature sensitivity of water absorption bands inthe glucose-measuring region can be a significant source of error inclinical assays. In addition, the devices can also be affected byindividual variations between patients at the measurement site. Skinlocation, temperature and tissue structure may affect the results, anddecrease the accuracy of the reading.

Other investigators have looked into measurement of glucose from bodyfluids other than blood, such as sweat; saliva, urine, or tears.However, factors relating to diet and exercise can affect glucose levelsin these fluids. In general, there is no strong correlation establishedbetween glucose concentration in the blood and in excreted fluids. Thelag time between blood and excreted fluid glucose concentrations can belarge enough to render such measurements inaccurate.

The continuous in vivo monitoring of glucose in diabetic subjects shouldgreatly improve the treatment and management of diabetes by reducing theonus on the patient to perform frequent glucose measurements. Implantedglucose sensors could be used to provide information on continuouslychanging glucose levels in the patient, enabling swift and appropriateaction to be taken. In addition, daily glucose concentrationmeasurements could be evaluated by a physician. An implantable sensorcould also provide an alarm for hypoglycemia, for example, overnight,which is a particular need for diabetics. Failure to respond can resultin loss of consciousness and in extreme cases convulsive seizures.Similarly, a hyperglycemic alarm would provide an early warning ofelevated blood glucose levels, thus allowing the patient to check bloodor urine for ketone bodies, and to avert further metaboliccomplications. (Jaffari, S. A. et al., “Recent Advances In AmperometricGlucose Biosensors For In Vivo Monitoring”, Physiol. Meas. 16:1–15(1995)).

Invasive glucose sensors may be categorized based on the physicalprinciple of the transducer being incorporated. Current transducertechnology includes electrochemical, piezoelectric, thermoelectric,acoustic, and optical transducers.

In piezoelectric, thermoelectric, and acoustic (surface acoustic wave,SAW) sensors used for glucose measurement, an enzyme-catalyzed reactionis used to create a measurable change in a physical parameter detectedby the transducer. The development of these sensors is at an earlylaboratory stage (Hall, E., Biosensors, Oxford University Press. Oxford,1990). Optical sensors are based on changes in some optical parameterdue to enzyme reactions or antibody-antigen reactions at the transducerinterface. Based on the nature of the monitoring process, they aredensitometric, refractometric, or colorimetric devices. At present, noneof them meets the selectivity requirements to sense and accuratelymeasure glucose in real physiological fluids.

There is a significant body of literature regarding the development ofelectrochemical glucose sensors. These generally incorporate an enzyme,which selectively reacts with glucose. Examples of enzymes, whichselectively react with glucose, are glucose oxidase (GOD), hexokinase,glucose-6-phosphate dehydrogenase (G-6-PD), or glucose dehydrogenase.Hexokinase is an enzyme that catalyzes the phosphorylation of glucose byATP to form glucose-6-phosphate and ADP.

Monitoring the reaction requires a second enzyme, glucose-6-phosphatedehydrogenase, in the following reaction:

The formation of NADPH may be measured by absorbance at 340 nm or byfluorescence at 456 nm (Jaffari, 1995).

Glucose dehydrogenase is another enzyme, which may be used formonitoring glucose in the following reaction:

The NADH generated is proportional to the glucose concentration.

Glucose oxidase is the most commonly used enzyme reported in theliterature. Its reaction is relatively simple, inexpensive, and may bemonitored using a variety of techniques.

These advantages have led to the extensive use of this enzyme inclinical analysis as well as its incorporation in the majority ofprototype biosensor configurations. The reaction of glucose with thisenzyme is a two-stage reaction:-D-glucose+GOD(FAD)→glucono-δ-lactone+GOD(FADH₂)  1)GOD(FADH₂)+O₂→GOD(FAD)+H₂O₂  2)glucono-δ-lactone+H₂O→gluconic acid  3)

The overall reaction is usually expressed as:β-D-glucose+O₂+H₂O→gluconic acid+H₂O₂  4)The reaction can therefore be monitored by the consumption of oxygen,the production of hydrogen peroxide, or the change in acidity due to theincrease of gluconic acid.

One of the key reasons for using these types of sensor in anintravascular environment, rather than subcutaneously or in other bodilyenvironments, is the need to provide closed-loop control for diabeticpatients. This would provide insulin delivery based on the patient'sactual glucose measurements, as opposed to providing insulin based onsome inexact approximation of the patient's glucose levels. This wouldof great benefit to diabetic patients. There is a widely recognized timedelay between glucose changes in venous blood, and subcutaneous glucosechanges. This time delay can range from just a few minutes, to up to 30min. However, the mathematical algorithm used to couple the glucosesignal to the insulin delivery system cannot tolerate a very long timedelay. In fact, two authors have presented data which suggested that 10minutes is the maximum delay which can be tolerated in closed-loopinsulin delivery systems (Parker R S, Doyle F, et al., “A Model-BasedAlgorithm for Blood Glucose Control in Type I Diabetic Patients” IEEETrans. Biomed. Engr. 46(2):148–157 (1999), and Gough D et al, “FrequencyCharacterization of Blood Glucose Dynamics” Ann Biomed Engr 31:91–97(2003).) Longer time delays can cause the controller to become unstable,potentially creating life-threatening issues for the patient, such asdelivery of extra insulin when blood glucose levels are falling rapidly.

Despite the foregoing and other efforts in the art, a suitablecontinuous in dwelling glucose sensor has not yet been developed.

A critical factor in the design of an implanted sensor is the anatomicalsite in which it is implanted. A few investigators have developedmonitoring systems, which can be placed within the vascular system.Armour et al. (“Application of Chronic Intravascular Blood GlucoseSensor in Dogs”, Diabetes 39:1519–26 (1990)) implanted a sensor into thesuperior vena cava of six dogs for a period of up to 15 weeks withrelative success. However, due to the risks of thrombosis andembolization, the majority of investigators have focused on subcutaneousimplantation.

A major drawback to subcutaneous implantation is the body's defenseagainst foreign objects: the “foreign-body response”. In this hostresponse, if an object cannot be removed by the inflammatory response,foreign-body giant cells will form a “wall” around the object, which issubsequently followed by the formation of a fibrous capsule. If theobject is a blood glucose sensor, it will no longer be in intimatecontact with body fluids, and the signal will drift and stability willbe lost. There are numerous reports of sensor stability being lost inabout a week (Wilson, G. S., et al., “Progress Towards The DevelopmentOf An Implantable Sensor For Glucose”, Clin. Chem. 1992 38:1613–7, andKerner, et al., “A Potentially Implantable Enzyme Electrode ForAmperometric Measurement Of Glucose”, Horm. Metab. Res. Suppl. Ser. 20:8–13 (1988)). Updike et al. (Updike, Stuart J., et al., “EnzymaticGlucose Sensors: Improved Long-Term Performance In Vitro And In, Vivo”,ASAIO J., 40: 157–163 (1994)) reported on the subcutaneous implantationof a sensor which was stable for up to 12 weeks, however, thisevaluation was only performed in three animals.

Recent clinical studies have also demonstrated that implantable insulinpumps are feasible for implantation for over one year (Jaremko, J. etal., “Advances Towards the Implantable Artificial Pancreas for Treatmentof Diabetes,” Diabetes Care, 21(3): 444–450 (1998)). The research wasinspired by the goal of the development of the artificial pancreas, andpromising initial clinical trials using implantable insulin pumps. Atthis point in time, development of implantable insulin pumps is at avery advanced stage, with units being implanted for over 2 years incanines (Scavani et al., “Long-Term Implantation Of A New ProgrammableImplantable Insulin Pump,” Artif. Organs, 16: 518–22 (1992)) and in 25patients for up to 3 years (Waxman, et al., “Implantable ProgrammableInsulin Pumps For The Treatment Of Diabetes”, Arch. Surg., 127: 1032–37(1992)).

A number of wearable insulin pumps are described by Irsigler et al.(“Controlled Drug Delivery In The Treatment Of Diabetes Mellitus,” Crit.Rev. Ther. Drug Carrier Syst., 1(3): 189–280 (1985)). Thus, it should berelatively straightforward to couple a long-term implantable glucosesensor as described in this disclosure, to an insulin pump to optimizeglycemic control for the patient.

In another aspect of this invention, it is possible to apply theprinciples discussed above to the direct, continuous monitoring ofarterial blood gases (ABG). Arterial blood gas values such as pO₂, pCO₂,and pH are the most frequently ordered laboratory examinations in theintensive care setting and the operating room (C. K. Mahutte,“Continuous intra-arterial blood gas monitoring,” Intensive Care Med(1994) 20:85–86). In the intensive care unit (ICU), ABG is typicallymonitored once a day, and additional measurements are only made once thepatient has experienced a deleterious event. Limited additional samplingis performed at the discretion of a physician or nurse. There can be asignificant time delay between the time the tests are ordered, and theresults are returned (E. E. Roupie, “Equipment Review: Continuousassessment of arterial blood gases,” Crit Care 1997 1(1):11–14).

In the case of continuous monitoring, significant changes in ABG valuesor trends would cause a rapid, therapeutic response on the part of thephysician, so potentially catastrophic events could be avoided.Continuous, non-invasive monitoring techniques such as pulse oximetryand continuous capnography have been introduced for this reason.Unfortunately, these devices are not always accurate in cases such asshock, hypothermia, or during the use of vasopressors. Further,pulse-oximetry does not measure oxygen tension.

A number of attempts have been made to develop improved arterial bloodgas monitors. There are two basic types of arterial blood gas monitors.In the first type, termed extra-arterial blood gas (EABG) monitors, thepatient's blood gas values are measured from a sample in the arterialcatheter. This can significantly reduce the time delay in obtainingresults, as compared to sending the sample to a laboratory. However, itis an on-demand system, not a continuous one, so the frequency ofsampling is once again dependent upon the physician or nurse.

The second type of device, known as an intra-arterial blood gas (IABG)monitor, is inserted directly into the arterial blood. However, theconsistency and reliability of these LABG monitors have not beenclinically acceptable because of problems associated with theintra-arterial environment. (C. K. Mahutte, “On-line Arterial Blood GasAnalysis with Optodes: Current Status,” Clin. Biochem 1998; 31:119–130).

In another aspect of this invention, implantable sensors capable ofmonitoring hemodynamic conditions over extended periods of time may beuseful in patients with heart failure. Such measurements havetraditionally been restricted to cardiac catheterization laboratoriesand intensive care units (ICU's). Measurements cannot be made easily inan ambulatory setting, or under conditions of cardiac loading, such asexercise. Ohlsson et al (Ohlsson A, et al., “Continuous ambulatorymonitoring of absolute right ventricular pressure and mixed venousoxygen saturation in patients with heart failure using an implantablehemodynamic monitor: Results of a one year multicentre feasibilitystudy,” Eur Heart J 22:942–954 (2001)) describe implantation of apressure sensor and an oxygen sensor, the IHM-1, Model 10040 (Medtronic,Inc.). However, in this study, 12 out of 21 oxygen sensors failed withinthe first 6 months of implantation. A fibrinous coating covered one ofthe sensors, and it was believed that this was responsible for sensorfailure. In addition, surgical implantation of this pacemaker-styledevice requires a 2–3 hour procedure in the operating room, which ismore costly than a catheterization procedure to insert a pressuresensor. Data from this study, however, points to the importance oflong-term, continuous pressure monitoring, as opposed to one-timemeasurements using standard cardiac catheterization techniques.

In another aspect of this invention, the sensor may be used formeasuring flow rates within a vessel. Congestive heart failure affectsmore than 5 million persons in the United States, and the rate isincreasing as people age and more of them survive heart attacks. Findingthe right treatment often involves a trial-and-error process, with thephysician trying different combinations of drugs and different dosagesto produce the best results. That means repeated blood pressure and flowmeasurements and invasive cardiac catheterization procedures to measurethe effects on the heart.

In U.S. Pat. No. 6,053,873 issued Apr. 25, 2000 to Govari et al, amethod of monitoring flow rates within a stent is described. Inparticular, the use of ultrasonic sensors to monitor blood flowrates isdescribed. However, due to the change in composition and thickness ofthe biological material present on the sensor surface, the boundaryconditions at the sensor may change over time. Initially, the sensorsurface, which may be calibrated against blood, may become covered witha layer of thrombus over time. Subsequently, the thrombus will transforminto fibrous tissue, affecting the absorption of the ultrasound signal.Blood may have an ultrasound absorption level similar to water, ofaround 0.002 dB/MHz cm, while the fibrous tissue layer may have anabsorption level similar to muscle, of around 2 dB/MHz cm. Thus, such asensor may require periodic re-calibration. Unfortunately, there-calibration process requires an invasive cardiac catheterizationprocedure, in which flow rates are determined using ultrasound,thermodilution, or the Fick method. In order to avoid additionalinterventional procedures, the fouling-resistant approach describedherein may be valuable.

In U.S. Pat. No. 6,309,350 to VanTassel, et al., issued Oct. 30, 2001, asensor is described which is anchored in the wall of the heart, or in ablood vessel. However, no mention is made of the difficulty with sensorfouling due to thrombus formation, or how to compensate for that.Pressure sensors would require recalibration as the softer thrombustransforms into neointimal tissue, or as the thickness of theencapsulating tissue layer changes. While the device could be initiallycalibrated during implantation, it would need to be recalibrated overtime. Because recalibration requires an invasive cardiac catheterizationprocedure, it would be desirable to avoid this if at all possible.

In addition, VanTassel et al. suggest the use of standard thermodilutionmethods to determine flow rates. Thermodilution is normally performed byinjection of either room temperature or iced saline through a catheter.However, it is desirable to avoid re-interventions if possible.Therefore, standard thermodilution methods are less desirable than thecurrent invention. In addition, it is impractical to use the sensor inthe standard method, i.e., by locally cooling the blood, becausechilling units are impractically large for implantation as a sensor.Further, thermodilution methods would be inaccurate if the thermocoupleor sensing element was covered by a relatively thick layer of cells.This would be especially true if only a small temperature rise wereintroduced into a large volume of flowing blood, such as in thepulmonary artery. Larger temperature rises (>2.5° C.) would cause localtissue damage.

Stroke is the third most common cause of death in the US and Europe.There are approximately 1 million acute ischemic strokes in Europeannually. Stroke survivors are often significantly disabled, and mustundergo extensive rehabilitation. The estimated short term costs of astroke are $13,649 per patient, and the long term costs are $45,893 fora minor stroke and $124,564 for a major stroke (Caro, J J, et al.,“Stroke Treatment Economic Model (STEM): Predicting Long-term Costs fromFunctional Status”, Stroke 1999; 30:2574–2579). The direct and indirecteconomic impact of stroke in the US is estimated to be $43.3 billion.(US). According to the National Institutes of Health recombinanttissue-type plasminogen activator trial, if stroke could be treated inthe first hours after the stroke occurs, the damage can be minimized(National Institute of Neurological Disorders and Stroke. rt-PA StrokeStudy Group. “Tissue Plasminogen Activator for acute ischemic stroke.” NEngl J Med 1995; 333:1581–1587). However, if treatment is delayed, thenthe damage becomes worse. Unfortunately, symptoms of stroke are notwidely recognized, and significant delays in treatment are introduced bythe failure to recognize the problem.

Risk factors such as high blood pressure, cholesterol, smoking, obesity;and diabetes increase the chances that an individual will suffer astroke. Patients with significantly increased risks of stroke includepatients with atrial fibrillation, coronary heart disease, asymptomaticcarotid stenosis, previous stroke, or transient ischemic attack (TIA).In the latter subsets, approximately 25% of stroke survivors experiencea recurrent stroke within five years, and among patients with TIA, therisk for stroke is 5% within 48 hours, 12% within 1 year, and up to 30%within 5 years.

In the acute phase of cerebral infarction, a great deal of experimentaldata suggests that free radicals, including superoxide, hydroxy radical,and nitric oxide (NO) are one of the most important factors causingbrain damage. J. Rodrigo, D, et al (Histol. Histopathol. 17, 973–1003(2002)) have observed that most of the morphological and molecularchanges associated with ischemic damage were prevented by treatment withinhibitors of NO production. In addition, there is a significantincrease in the concentration of NO following cerebral ischemia. Thereis also significant NO release in hemorrhagic stroke. (Chen H H, et al,“Low cholesterol in erythrocyte membranes and high lipoperoxides inerythrocytes are the potential risk factors for cerebral hemorrhagicstroke in human” Biomed Environ Sci. 2001 September; 14(3):189–98).Therefore, the detection of NO in the acute phase of ischemic orhemorrhagic stroke would provide an early detection method for stroke,allowing treatment to be performed as promptly as possible by thephysician or health-care provider.

While animal studies have demonstrated an increase in nitric oxideshortly after ligation of the middle cerebral artery (Lin, S Z, et al.,“Ketamine Antagonizes Nitric Oxide Release From Cerebral Cortex afterMiddle Cerebral Artery Ligation in Rats”, Stroke 1996; 27:747–752) it ispossible that the nitric oxide may be converted to metabolites such asnitrites and nitrates by the time it reaches monitoring sites which aredistant from the site of the stroke. It has also been demonstrated thatthe level of these metabolites are significantly higher in cerebrospinalfluid (CSF) at the time stroke patients are admitted to the hospital,which is often many hours after the stroke (Castillo, J. et al., “NitricOxide-Related Brain Damage in Acute Ischemic Stroke”, Stroke 2000;31:852–857). It has further been demonstrated that there is asignificant increase in levels of nitric oxide, nitrite, and nitrateions in the jugular vein of a rat, immediately following induction ofischemic stroke (Suzuki, M, et al., Brain Research 951 (2002) 113–120).

Notwithstanding the extensive efforts in the prior art, however, thereremains a need for an implantable blood glucose sensor for implantationin a blood vessel, which can provide useful blood glucose readings foran extended period of time, without material interference from thrombusformation, embolization, or other foreign body response. Preferably, thesensor is capable of continuous or near continuous monitoring, anddriving an implantable insulin pump and/or making blood glucose dataavailable to the patient or medical personnel.

SUMMARY OF THE INVENTION

The present invention generally relates to the use of sensors to monitorthe concentration of a chemical species in bodily fluids, and morespecifically, to a novel sensor configuration to monitor glucose levelsin a body vessel. The device is an implantable sensor, which isdelivered to the patient's vascular system preferably transluminally viaa catheter, using a stent or stent-graft as a platform. One feature ofthe device is that the sensor surface is placed at the apex of theluminal surface of a streamlined housing, so that the shear rate at thesensor/blood interface is sufficient to minimize the thickness of theformed thrombus layer. In this manner, significant tissue deposition orencapsulation due to potential fibrotic reactions is minimized, andtransport of glucose to the sensor is not altered over time.

Thus, there is provided in accordance with one aspect of the presentinvention a blood glucose detector for implantation within a bloodvessel. The blood glucose detector comprises a support, having a firstside for contacting the wall of the vessel and a second side for facingradially inwardly towards the center of the vessel. A sensor is carriedby the support, and the sensor has a sensing surface thereon. Thesensing surface is spaced radially inwardly from the first side by adistance of at least about 0.2 to 2.5 mm, such that the velocity ofblood in the vessel inhibits obstruction of the sensing surface.Preferably, the distance is at least about 0.5 mm. The blood glucosedetector further comprises a transmitter on the support, fortransmitting information from the sensor to an external receiver. In oneembodiment, the support comprises an expandable tubular body. Thetubular body may be either a balloon expandable or a self-expandablecomponent such as a stent. The tubular body may be further provided witha tubular sheath on the radially inwardly directed surface and/or theradially outwardly directed surface. In one embodiment, the sensorcomprises an analyte permeable membrane and an enzyme gel layer.

In accordance with another aspect of the present invention, there isprovided a method of prolonging the useful life of a sensor in a bloodvessel. The method comprises the steps of providing a sensor having ananalyte sensing surface thereon, and positioning the sensor at a site ina blood vessel such that the sensing surface is positioned radiallyinwardly from the vessel wall by a sufficient distance that the bloodflow shear rate at the sensing surface substantially delays obstructionof the sensing surface. Preferably, the positioning step comprisescarrying the sensor on a catheter and transluminally advancing thecatheter to the site.

In accordance with a further aspect of the present invention, there isprovided an implantable sensor for sensing the presence of an analyte ina vessel. The sensor comprises a tubular support structure for anchoringthe sensor in a vessel. The support has a sidewall with a luminal sidefacing towards the center of the vessel and an abluminal side facingtowards the wall of the vessel. A sensor housing is carried by thesupport structure, the housing having a streamlined exteriorconfiguration to minimize blood flow turbulence. A power supply andelectrical circuitry are provided in the housing, and a sensing surfacecarried by the housing is exposed to the exterior of the housing. Thesensing surface is positioned on the radially inwardly most portion ofthe luminal side of the housing.

In accordance with a further aspect of the present invention, there isprovided an implantable sensor for sensing the presence of an analyte ina vessel that can be retrieved. The sensor comprises a support structurefor anchoring the sensor in a vessel. The sensor further comprises asnareable member connected to the sensor that allows allow for removalof the sensor in a catherization procedure. In one embodiment, thesnareable member is a hook.

In accordance with a further aspect of the present invention, a methodfor retrieving an implanted sensor is provided. Under fluoroscopicguidance, a guiding catheter of sufficient diameter so as to be able toaccommodate the retrieved sensor and its anchoring platform is inserted.A snare is inserted through the guiding catheter and is guided to asensor hook. The snare grasps the sensor hook and the sensor collapsesinto a retrieval catheter. The sensor and guiding catheter aresimultaneously withdrawn from the patient's body.

In accordance with a further aspect of the present invention, anothermethod for retrieving an implantable sensor on a support is provided.Under fluoroscopic guidance, a catheter with a clip is positioned sothat the clip is adjacent to an sensor attached to a stent. A balloonattached to the catheter is inflated so that the clip is forced aroundthe sensor. After deflating that balloon, another balloon is inflated sothat the clip pulls the sensor away from the stent, thus separating thesensor therefrom.

In accordance with a further aspect of the present invention, there isprovided an implantable immunosensor. The immunosensor comprises asupport structure. The immunosensor produces an electrical signalrepresentative of a reaction between an analyte and an antigen.

In another aspect of this invention, the sensor may be an infraredsensor placed either intraluminally or extraluminally in a blood vessel.The infrared sensor produces an electrical signal indicative of theconcentration of chemical compounds in blood.

In accordance with a further aspect of the present invention, there isprovided an implantable blood gas monitor. The blood gas monitorproduces an electrical signal indicative of the partial pressure ofdissolved blood gases or pH.

In accordance with a further aspect of the present invention, there isprovided an implantable ion-selective electrode. The ion selectiveelectrode produces an electrical signal indicative of the concentrationof electrolytes in blood.

In accordance with a further aspect of the present invention, there isprovided an implantable pressure sensor. The pressure sensor produces anelectrical signal indicative of the pressure in the vessel.

In accordance with a further aspect of the present invention, there isprovided an implantable flow sensor. The flow sensor produces anelectrical signal indicative of the flowrate in the vessel.

In accordance with a further aspect of the present invention, sensor forimplantation within a blood vessel that minimizes the formation ofthrombus. The sensor comprises a support, having a first side forcontacting the wall of the vessel and a second side for facing radiallyinwardly toward the center of the vessel and a sensor carried by thesupport and having a sensing surface thereon. The sensing surface of thesensor is spaced radially inwardly from the first side and includes alayer that minimizes the formation of thrombus. The sensor can include atransmitter on the support, for transmitting information from the sensorto an external receiver. The layer can be an anticoagulant (such asheparin), a hydrogel (such as poly(ethylene glycol), poly(N-vinylpyrrolidone), or poly(hydroxyethylmethacrylate), or can release apharmacological agent that inhibits cell proliferation or migration.

In accordance with a further aspect of the present invention, there isprovided an implantable sensor for sensing the presence of nitric oxideor a nitric oxide metabolite in a vessel. The sensor comprising asupport structure, a sensor housing carried by the support structure,and a sensing surface exposed to the surrounding environment. The sensorcan detect nitric oxide or a nitric oxide metabolite and can beimplanted for a period of at least one week. The support structure canbe a stent or a catheter.

Further features and advantages of the present invention will becomeapparent to those of skill in the art in view of the detaileddescription of preferred embodiments which follows, when consideredtogether with the attached drawings and claims.

BRIEF DESCRIPTIONS OF THE DRAWINGS

FIG. 1A is a perspective view of an expanded stent with an embeddedsensor housing on its abluminal side.

FIG. 1B is a block diagram of remote circuitry for an externalmonitoring device.

FIG. 1C is a diagram of a wearable or implantable insulin pump.

FIG. 1D is a block diagram of the sensor circuitry for measuring analyteconcentration.

FIG. 1E is a block diagram of the remote measurement unit.

FIG. 2 is a perspective partial cut away view of a stent sensor devicesurrounded by a sheath.

FIG. 3 is a cross-section taken along the line 3—3 in FIG. 2.

FIG. 4 is an enlarged cross-sectional view through the sensor element ofFIG. 3.

FIG. 5A is a cross-sectional view through a stent showing one mountingconfiguration of an embedded sensor in accordance with one embodiment ofthe present invention.

FIG. 5B is a cross-sectional view as in FIG. 5A of a stent with analternate configuration for the embedded sensor.

FIG. 5C is a cross-sectional view as in FIG. 5A of a stent with anembedded sensor mounted completely on the luminal side of the stent.

FIG. 6A is a perspective view of an alternate support structure inaccordance with the present invention.

FIG. 6B is a perspective view of a further support structure inaccordance with the present invention.

FIG. 6C is a perspective view of the support structure illustrated inFIG. 6A, provided with an outer fabric sheath.

FIG. 7A is a side elevational view of the distal end of a catheter-baseddelivery system for a stent sensor device.

FIG. 7B is a cross-sectional view of a catheter-based delivery systemfor a rolled sheet type self-expanding stent with an embedded sensor.

FIG. 7C is a cross-sectional view of a catheter-based delivery system,which uses a restraining sheath, for a self-expanding stent with anembedded sensor.

FIG. 7D is a cross-sectional view of a catheter-based delivery system,with a pull-away sheath, for a self-expanding stent with an embeddedsensor.

FIG. 7E is a cross-sectional view of a catheter-based delivery systemfor a balloon expandable stent with an embedded sensor.

FIG. 7F is a cross-sectional view of a catheter-based delivery systemfor a self-expanding stent with an embedded sensor, in which theguidewire lumen passes on either side of the sensor.

FIG. 8 is a cross-sectional view of the catheter-based delivery systemtaken along the 8—8 line of FIG. 7A.

FIG. 9A is a cross-sectional view through a stent showing a transduceracross the cross-section of the stent in accordance with the presentinvention.

FIG. 9B is a perspective view of an expanded stent with an embeddedsensor housing on its abluminal side and a transducer across thecross-section of the stent.

FIG. 9C is an enlarged cross-sectional view of the transducer of FIG.9B, taken from the 9C—9C line.

FIG. 9D is a perspective partial cut away view of a stent sensor devicesurrounded by a sheath with a transducer partially across thecross-section of the stent.

FIG. 9E is a perspective view of an expanded stent with an embeddedsensor housing on its abluminal side and four perpendicular transducersplaced partially across the cross-section of the stent.

FIG. 10 shows a side elevational view of a sensor and transmitter, withexpanded anchoring stents at the proximal end of the sensor,intermediate between the sensor and transmitter, and at the distal endof the transmitter.

FIG. 11 shows a side elevational view of a sensor and transmitter, withanchoring stents in the compressed state at the proximal end of thesensor, intermediate between the sensor and transmitter, and at thedistal end of the transmitter.

FIG. 12 is a diagram of a constrained anchoring platform or stent with asensor housing containing a reservoir for delivery of labeled targetmolecules, and a sensing compartment containing a signal source, areceiver and antibodies with and without labeled target molecules,signal processing circuitry, and a radiotransmitter.

FIG. 13 is a block diagram of the fluorescence based sensor electronicsof the sensor of FIG. 12.

FIG. 14A is a diagram of an expanded anchoring platform or stent with anembedded sensor housing on its luminal side, and containing a hook withwhich the device can be retrieved.

FIG. 14B is a diagram of an expanded anchoring platform or stent with anembedded sensor housing on its luminal side, and containing a hook onthe sensor housing with which the device can be retrieved.

FIG. 15A is a side view diagram of an expanded anchoring platform withan embedded sensor, which is placed centrally in a vessel, andcontaining a hook on the sensor housing with which the device can beretrieved.

FIG. 15B is an end view diagram of an expanded anchoring platform withan embedded sensor, which is placed centrally in a vessel, andcontaining a hook on the sensor housing with which the device can beretrieved.

FIG. 16 is a side view diagram of a catheter containing a sensor nearits distal end, with self-expanding split-ring anchors.

FIG. 17A shows a side view of a retrieval catheter designed for removinga sensor at the end of its life.

FIG. 17B shows an end view of a retrieval catheter designed for removinga sensor at the end of its life.

FIG. 18 shows a side view of a catheter containing a nitric oxidesensor.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT

In accordance with the present invention, an intraluminal blood glucosesensor is provided on a support structure such as a modified stent ofthe type implanted following percutaneous transluminal coronaryangioplasty (PTCA) for the treatment of atherosclerosis. Atherosclerosisis the build-up of fatty deposits or plaque on the inner walls of apatient's arteries. These lesions decrease the effective size of theartery lumen and limit blood flow through the artery, prospectivelycausing a myocardial infarction or heart attack if the lesions occur incoronary arteries that supply oxygenated blood to the heart muscles. Inthe angioplasty procedure, a guide wire is inserted into the femoralartery and is passed through the aorta into the diseased coronaryartery. A catheter having a balloon attached to its distal end isadvanced along the guide wire to a point where the stenotic lesionslimit blood flow through the coronary artery. The balloon is theninflated, compressing the lesions radially outward against the wall ofthe artery and substantially increasing the size of the arterial lumen,to improve blood circulation.

Increasingly, stents are being used in place of or in addition to PTCAfor treatment of atherosclerosis, with the intent of minimizing the needto repeatedly open a stenosed artery. Although a number of differentdesigns for stents exist in the prior art, all are generally configuredas elongate cylindrical structures that can assume two different states,one having a substantially greater diameter than the other. A stent isimplanted in a patient's vascular system using an appropriate deliverysystem. There are two basic types of stents. The first type is termed“balloon expandable”, and refers to sterits that are expanded radiallyoutward due to the force from an inflated angioplasty balloon, such asthe Palmaz-Schatz stent, the Gianturco-Roubin stent, and the Streckerstent. The second type is termed “self-expandable”, and refers to andthose that are self-expanding, such as the Hemobahn™ and SMART Stent™(made of nickel titanium memory alloy), and the Wallstent (made ofElgiloy).

Typically, a stent carried by a delivery catheter is advanced through aguide catheter to a site within the patient's artery. For the balloonexpanded type, after the introducer sheath is retracted, a balloondisposed inside the stent is inflated to a pressure ranging from aboutsix to ten atmospheres. The force produced by the inflated balloonexpands the stent radially outward beyond its elastic limit, stretchingthe vessel and compressing the lesion to the inner wall of the vessel. Aself-expanding stent expands due to spring force following itspositioning within the artery, after a restraining sheath is retractedfrom the compressed stent. Following the expansion process, if theballoon expandable type is used, the balloon is deflated and removedfrom inside the stent and the catheter and other delivery apparatus iswithdrawn. The lumen through the vessel should be substantiallyincreased, improving blood flow.

After a stent or other endoluminal device is implanted, a clinicalexamination and either an angiographic or ultrasonographic morphologicalprocedure is performed to evaluate the success of the procedure inopening the diseased artery or vessel. These tests are typicallyrepeated periodically, e.g., at six-month intervals, because restenosisof the artery may occur.

Although the sensor of the present invention may be carried by any of awide variety of intraluminal support structures, balloon expandable orself-expandable stents are preferred by the present inventor. Ingeneral, the stent may be modified from those intended for supporting atreatment site following angioplasty, in view of the preferred placementin a healthy section of the vessel as is discussed in greater detailbelow. In addition, the wall of the stent may be modified to support thesensor, as is also discussed in greater detail below. As well asproviding a useful support structure with known deploymentcharacteristics, a stent provides a useful platform for a variety ofadditional reasons. For example, it is impractical to pass an electronicconductor through the wall of an artery or vessel to monitor thecondition of an implanted sensor for long periods of time. Also, anyactive glucose sensor would likely be energized with electrical power.Again, it is less practical to supply power to such a sensor through anyconductor that perforates the vessel wall or that passes outside thepatient's body.

In addition to stents, the generic term endoluminal implant encompassesstent-grafts, which are also sometime referred to as “covered stents.” Astent-graft is a combination of one or more stents and a synthetic graftthat is typically implanted at a desired point in a vessel using anendoscopic approach. The stent is attached either to the ends orthroughout the body of the synthetic graft, and is used to hold thegraft in position. Sometimes, hooks are provided on the stent to ensurethat the graft remains in the desired position within the vessel.Clearly, it would also be desirable to monitor the status of glucose andother parameters through a stent graft, just as noted above in regard toa stent.

Endoluminal implants are used in other body passages in addition toblood vessels. For example, they are sometimes used to maintain an openlumen through the urethra, or through the cervix. A stent placedadjacent to an enlarged prostate gland can prevent the prostate fromblocking the flow of urine through the urinary tract. Tracheal andesophageal implants are further examples of endoluminal implants. Inthese and other uses of an endoluminal implant, provision for monitoringparameters related to the status of flow and other conditions in thepatient's body would be desirable. Information provided by monitoringsuch parameters can enable more effective medical treatment of apatient.

A number of factors will affect the function of a glucose sensorimplanted within a blood vessel. As with stents, the device should bedesigned to minimize the risk of thrombosis and embolization. Thus,slowing or stoppage of blood flow at any point within the lumen shouldbe minimized. Colombo et al. (“Intracoronary Stenting WithoutAnticoagulation Accomplished With Intravascular Ultrasound Guidance,”Circulation 91:1676–88, (1995)) and Goldberg et al. (“Benefit OfIntracoronary Ultrasound In The Deployment Of Palmaz-Schatz Stents”, J.Am. Coll. Card. 24: 996–1003, (1994)) demonstrated that a factorresulting in subacute stent thrombosis in coronary arteries wasinadequate expansion of the stent. This was done using intravascularultrasonography (IVUS). Another factor, and related, is that the stentsshould be placed in close apposition to the vessel wall, so that theonly flow obstructions are at the luminal surface of the stent struts.They later demonstrated that by using non-compliant balloons at highpressure (18 atm), that even the final assessment using IVUS is notrequired; an important consideration for centers lacking such equipment.Thus, it is preferable to minimize flow disturbances which may becreated by stent implantation. This is more preferable in the smallercoronary arteries, which range from about 2.5–4 mm in diameter, than itis in other, larger arteries, where small amount of mural thrombus willhave less effect on the volumetric flowrate.

Another factor is that the stent platform is not a solid piece ofimpermeable material, such as a rolled sheet, or a “jelly-roll”, asdescribed by Winston et al. (U.S. Pat. No. 5,411,551 issued May 2,1995). In in vivo studies, Virmani, et al. (“Histopathologic EvaluationOf An Expanded Polytetrafluoroethylene Nitinol Stent Endoprosthesis InCanine Iliofemoral Arteries,” JVIR, 10:445–456 (1999)) demonstrated thatwhen endovascular stent-grafts are covered with porous graft material,it is possible for the graft material, in this case ePTFE, to becomecovered with a nearly intact layer of endothelial cells. However, whenthe same graft material was rendered impermeable to the underlying hostvessel by wrapping it in FEP, the graft material was instead coveredwith a thick layer of thrombus. This demonstrates the desirability ofusing porous stents as anchors for the sensor.

Usually, a layer of fibrin clot deposits on the luminal surface of theimplanted stent. After 3–4 weeks, the fibrin layer is typicallyremodeled into fibro-collagenous tissue, with a layer of endothelialcells present on the flow surface. However, if a relatively thick layerof fibro-collagenous tissue forms on the surface of an endovascularsensor, it may experience the same loss of signal that has been reportedfor subcutaneously implanted sensors. That is, a modified type of“foreign-body response” may occur, and glucose may not be able to freelyand rapidly diffuse to the sensor. Evaluation of an ePTFE (expandedpolytetrafluoroethylene) covered stent graft (Virmani, et al., 1999)showed that in areas where the graft surface was further from the centerof the flow channel, it became covered with a relatively thick layer ofthrombus, which eventually converted to fibro-collagenous tissue. Inareas where the graft surface was closer to the center of the flowchannel, it became covered with a relatively thin layer of tissue, andin some cases remained uncovered for periods of up to one year in acanine. Thus, there is a balance between minimizing disruption of bloodflow, and excessive thrombus deposition.

Unlike a stent or stent-graft, which is used in the treatment ofocclusive or aneurysmal vascular disease, however, the stent-sensorcombination of the present invention should be placed in a relativelyhealthy portion of the artery. If it is placed in a stenotic orcalcified area, it may be difficult for the device to expand completely,so the device may be more narrowed than is intended, resulting inthrombotic fouling of the sensor, or distal embolization may occur. Ifthe device is placed in an aneursymal vessel, it may expand more thanintended, and again, fail to function as designed.

In addition, some types of sensors require that the actual sensingcomponent be covered with a semi-permeable membrane in order to allowcertain components such as glucose to pass through, while excludingcells, enzymes, or other factors which could damage the sensor. Thus, astent-graft-sensor combination might be more suitable than astent-sensor combination without the graft material. Furtherconsiderations will become apparent from the illustrated embodiments ofthe present invention which are discussed below.

FIG. 1A shows a perspective view of an expanded implantable sensordevice 10 having a proximal end 12, a distal end 5 and a central lumen16 extending therethrough. The expanded implantable sensor device 10 iscomprised of a stent 14 and a sensor 20. Although the stent 14illustrated in FIG. 1A resembles a Palmaz-Schatz stent, it is understoodthat any of a wide variety of stent configurations can be utilized suchas those identified below. Whether the stent is a balloon expandable orself-expandable variety, the stent 14 comprises a tubul sidewall 18,which may be compressed to a relatively small diameter for deliveryusing a catheter, and which may be expandable to the approximatediameter of the vessel or to a diameter approximately 5–25% greater thanthe diameter of the vessel lumen.

In general, a self-expanding variety of stent or stent-graft such as theS.M.A.R.T. Stent (Cordis Corp., Miami, Fla.) or Hemobahn (W.L. Gore &Associates, Flagstaff, Ariz.) is preferred, because the presence of thesensor body may impede proper expansion of the stent platform using aballoon. The sensor housing can preferably be attached to the stent byuse of adhesives, by tying (suturing) the two components together, or byuse of thermal bonding, soldering, welding or brazing, or mechanicalinterfit or the like. The sensor and sensor housing should be attachedto the stent along a single row of stent struts, so that the struts mayfully expand without needing to stretch the sensor housing, or beingconstrained by the sensor housing.

An example of the type of stent which could be used in certainembodiments of the present invention is described by Lau et al. in U.S.Pat. No. 5,876,432, issued Mar. 2, 1999 which is herein incorporated byreference. Other stents which are suitable for use in the presentinvention are described by U.S. Pat. No. 4,739,762 to Palmaz, issuedApr. 26, 1988; U.S. Pat. No. 5,102,417 to Palmaz, issued Apr. 7, 1992;U.S. Pat. No. 5,421,955 to Lau et al., issued Jun. 6, 1995; U.S. Pat.No. 5,195,984 to Schatz, issued on Mar. 23, 1993; U.S. Pat. No.4,886,062 to Wiktor , issued Dec. 12, 1989; U.S. Pat. No. 4,655,771 toWallsten, issued Apr. 7, 1987; U.S. Pat. No. 5443,500 to Sigwart, issuedAug. 22, 1995; and U.S. Pat. No. 4,580,568 to Gianturco, issued Apr. 8,1986; which are all incorporated in their entireties herein byreference. The expanded inside diameter of the device may range from 3to 20 mm, and the length of the device may range from 1 to 30 cm. Theradially inwardly facing or radially outwardly facing sidewall of thestent may also be covered by a sheath as is known in the art. The sheathcan be bonded to the surface of the stent so that the sheath remains inclose apposition to the vessel wall.

Attached to the stent 14, a sensor 20 contains the sensing circuitrywithin its housing. Depending upon the stent 14 design and sensor 20design, the attachment between the sensor 20 and the stent can take anyof a variety of forms as will be apparent to those of skill in the artin view of the disclosure herein. For example, adjacent axiallyextending parallel struts of the stent design illustrated in FIG. 1Aspread circumferentially apart upon expansion. Thus, the sensor 20should be mounted to only colinear axial struts along the length of thestent unless the mounting is designed to accommodate circumferentialseparation of parallel struts. Alternatively, the sensor 20 may bepositioned in an aperture which is cut in the sidewall of the stent 14,and attached to the strut portions adjacent to the edge of the aperture.Attachment may be accomplished in any of a variety of ways, such as bythe use of adhesives, thermal bonding, soldering, welding or brazing, ormechanical interfit, depending upon the construction materials of thestent 14 and the sensor 20. Alternatively, the sensor 20 may be solventbonded, adhesively bonded or otherwise attached to an expandable tubularsleeve which is positioned concentrically about the stent 14. Otherattachment configurations can be devised by those of skill in the art inview of the disclosure herein.

In the illustrated embodiment, all or a portion of the sensor 20 ispositioned on the radially outwardly facing surface of the tubularsidewall 18. Positioning the sensor 20 on the radially outwardly facingsurface (sometimes referred to herein as the abluminal surface) mayadvantageously optimize blood flow through the central lumen 16 and/orreduce turbulence in the central lumen 16. Alternatively, the sensor 20may be positioned on the radially inwardly facing surface (sometimesreferred to herein as the luminal surface) of the stent. In addition,the stent struts or other elements which make up the sidewall 18 may beremoved or modified at the portion of the sidewall corresponding to thesensor 20 such that the sensor 20 may be mounted within the resultingopening in the sidewall. Because the implantable sensor device 10, inaccordance with the present invention, is preferably positioned within ahealthy section of the vessel, the radial strength normally required fora stent in a post-angioplasty application is unnecessary. Thus, thepresent inventor believes that any reduction in radial support which mayresult from reasonably dimensioned apertures or other modifications tothe sidewall to attach sensor 20 will not adversely affect the role ofthe stent as a sensor support structure in the context of the presentinvention. Additional details concerning the position of the sensor 20with respect to the tubular wall 18 will be discussed below inconnection with FIGS. 5A–5C.

The sensor circuitry 22 includes a sensing circuit 24, which isconnected to a signal processing circuit 26, which is connected to apower source 28, a radio transmitter 30, and an antenna (not shown) totransmit signals about the glucose concentration to a remote device. Theantenna (not shown) may be a thin wire wound around the stent 10, orpreferably, may be wound around a small ferrite core, as commonly usedfor such applications. This provides certain advantages, such asincreased signal strength or increased transmission distance for thesame amount of power required. In Keilman et al. (U.S. Pat. No.6,231,516, issued May 15, 2001), for example, the implied requirement tominimize protrusion of the sensor and transmitter into the vessel lumenrequires that the stent be constructed using unwieldy manufacturingprocesses, and the signal strength is weakened by the inability to use aferrite core.

The power source 28 may be inductively coupled to an external device, orit may be a thin film rechargeable battery, as in Bates, J. B. et al.,“Thin Film Rechargeable Lithium Batteries for Implantable Devices”,ASAIO J., 1997 43:M644–M647, which is incorporated herein by reference.Alternatively, the power source may be a battery, such as lithiumiodide, lithium silver vanadium oxide, lithium carbon monofluoride, orlithium ion rechargeable battery, as commercially available from WilsonGreatbatch Technologies, Inc. (Clarence, N.Y.).

The glucose sensor monitoring circuit and the remote measurement unitcircuit are shown in FIGS. 1D and 1E. The glucose sensor monitoringcircuit consists of a voltage source (25) to drive the sensor (20), anda current-to-frequency converter (27). The circuit is powered by anexternal field produced by the remote measurement device. In normaloperation, the remote measurement device (FIG. 1E) is placed over thearea where the sensor is implanted. This produces a voltage across thePower Coil and Antenna (30) that is then regulated by the power supply(28) to provide a source of power for the sensor electronics (voltagesource and current-to-frequency converter). Once powered, the circuitwill produce an output frequency that is directly proportional to theconcentration of glucose.

FIG. 1E shows a block diagram of the electronics for the remotemeasuring unit. The remote monitoring unit provides the drive to powerthe sensor electronics and also receives and processes the signal fromthe V-F converter. A high current drive circuit (37), running at afrequency much lower than the V-F converter, is used to excite theantenna/excitation coil (34). Because the signal received from thesensor is much higher than the excitation signal, the former is filteredand amplified in amplifer/filter (35). The microcontroller (39) measuresthe frequency and then uses a calibrated look-up table to providetranslation to the proper units and also to compensate for anynon-linearities in the device response curves. The final result isdisplayed on an LCD (not shown) for the user.

The sensing circuit 24 is in electrical communication with a sensingsurface (not illustrated in FIG. 1A) for sensing the analyte of interestin the fluid stream. In the embodiment illustrated in FIG. 1A, thesensor surface is preferably positioned radially inwardly from theradially inwardly facing surface of the tubular sidewall 18, to improvethe useful life of the device as is discussed elsewhere herein. Thus, inan embodiment such as that illustrated in FIG. 1A in which the sensor 20is positioned on the abluminal side of the tubular sidewall 18, thesensor surface is displaced radially inwardly such as by a radiallyinwardly extending portion of the housing or other support (notillustrated) to achieve the desired radial position of the sensorsurface. Alternatively, in an embodiment in which the sensor 20 ispositioned on the luminal side of the tubular side wall 18, or within orthrough an opening in the tubular sidewall 18, the sensor surface may bepositioned directly on a radially inwardly-most extending portion of thesensor 20 as is discussed elsewhere herein.

The sensor surface is preferably covered by a semi-permeable membrane(not shown), which contacts passing blood when the stent 14 is placed ina blood vessel. The permeability of the membrane is selected to allowblood glucose, or the analyte of interest to freely contact the sensor,while restricting the passage of other blood components. Thesemi-permeable membrane may comprise ePTFE, Dacron®, polyurethane,silicone rubber, poly(lactide-co-glycolide) (PLGA), poly(caprolactone)(PCL), poly(ethylene glycol) (PEG), collagen, polypropylene, celluloseacetate, poly(vinylidene fluoride) (PVDF), nafion or other biocompatiblematerial. These membrane materials may also be used for the tubularsheath which can be used to surround either the luminal or abluminalsurfaces of the stent. The membrane of the sheath may be bonded to thestent by adhesives or by tying (suturing) the two components together orby use of thermal bonding, soldering, welding or brazing, or mechanicalinterfit. The pore size of the lumenal membrane may be large enough toallow cells to come through if it covers the sensor surface, but thesensor membrane should have a pore size (MW cutoff of about 1000) whichwill allow glucose to pass, while restricting the passage of largermolecules. The entire tube surface does not have to be composed of thesame membrane material. The part of the device near the sensing elementcan be composed of a different material or have a different porositythan the material in the rest of the device.

Referring to FIG. 1B, a remote circuit 32 is equipped with an antenna 34and a signal processing unit 36, which converts the electronic signalfrom the embedded sensor into a concentration level or other indicium ofthe analyte. Preferably, an alarm circuit 38 and a display 40 are alsoprovided. Information regarding the level of the analyte of interest canbe displayed on the display 40 such as a monitor 42. The signalprocessing unit 36 may be provided with a lookup table or other baselineof normal or expected values for the analyte of interest. If theconcentration of the analyte goes outside of the prescribed range, or ifan electronic failure is detected, a warning audible, visible, ortactile signal is preferably produced from the alarm system 38. Atransmitter 44 may also be included in the remote circuitry 32 in orderto transmit data about the level of the analyte of interest to animplantable infusion pump, or the like.

The remote circuitry 32 can be provided in any of a variety of physicalforms, depending upon the intended use of the device. For example, in ahospital or other immobilized patient setting, the remote circuitry 32can be provided in a desktop or bedside housing and coupled directly toa display 40 such as a monitor 42. Alternatively, ambulatory patientdevices may be provided by deleting a permanent coupling to the monitor42 and packaging the remaining components of remote circuit 32 in awearable form, such as a compact self-contained unit adapted forattachment to the wearer's clothing or including straps so that it canbe strapped to the patient's body. In an ambulatory device, the signalprocessing unit 36 includes sufficient memory to record the glucosevalues over a predetermined period of time. The patient can couple thewearable device to an external monitor 42 or other electronicsperiodically, such as one or more times per day. Analyte data maythereafter be displayed in readable form, such as on a monitor 42 intable form or in graph form with a time unit on the X axis and a glucosevalue or derivative data on the Y axis.

The wearable unit (not illustrated) may additionally be provided with adata export coupling, such as a telephone connector, for connecting thesignal processing unit 36 via internal or external modem into thetelephone system. In this manner, the patient can transmit condensedanalyte data to the healthcare provider for further monitoring and/oranalysis. The wearable unit may additionally be provided with one ormore manual data inputting elements ranging from a simple push button toa keypad, to allow manual data entry relating to significant dietary orother events. For example, meal times, significant fluid intake, manualinsulin injection or other administration, or any of a variety of othersignificant events can be marked in the data, so that the patient orreviewing medical personnel can correlate the event with the bloodglucose data.

Referring to FIG. 1C, there is schematically illustrated an implantableor externally wearable infusion pump 46. The infusion pump 46 may becontrolled by the remote circuit 32 via a receiver 48 in the pump, ormay be manually controlled using sensor information only as a guideline.The infusion pump 46 may be refilled through an appropriately designedport 50 such as a pierceable septum using a hypodermic needle or otherappropriate delivery system. The infusion pump 46 may be implantable, asdescribed by Irsigler et al., (“Controlled Drug Delivery in theTreatment of Diabetes Mellitus”, Crit. Rev. Ther. Drug Carrier Syst.1(3): 189–280 (1985)), which is incorporated herein by reference.Alternatively, it may be worn externally by the patient, and infuseinsulin or other drugs as appropriate through a catheter 15 which isinserted into the patient's body. External insulin infusion pumps arecurrently marketed by suppliers like Medtronic, or Siemens. However,these pumps are not designed to receive a continuous signal from animplanted sensor, but instead are pre-programmed to approximate thepatient's baseline insulin requirements. Such pumps can be modified withan with appropriate circuitry to receive and respond to output from theglucose sensor by those of skill in the art in view of the disclosureherein.

Now referring to FIG. 2, a covered implantable sensor device 10 isshown. The sensor 10 comprises a cylindrical stent wall 18 surrounded bya sheath 52. An antenna (not shown) may be wound around the body of thesensor 10 and connected to the power source or the transmitter. Allrelevant electronics are schematically illustrated as in electronicshousing 54 which is electrically coupled to a sensor 56 by one or moreconductors 57. All such junctions of dissimilar metals are coated withpolymers which are impermeable to bodily fluids in order to reducegalvanic corrosion. The analyte sensing element 56 is covered with amembrane 62 (FIG. 4) which is permeable to the analyte of interest. Theanalyte sensing element 56 extends radially inwardly within the sensor10 where blood flow conditions are optimal.

Now referring to FIGS. 2, 4 and 5, the illustrated analyte sensingelement 56 contains an enzyme gel layer 64, which is placed adjacent tothe outer permeable membrane 62. The analyte diffuses through themembrane 62 to the gel enzyme layer 64. The reaction between the analyteand the enzyme occur in the gel enzyme layer 64. The reaction productsthen pass through an inner membrane 66 and react at the surface of anoble metal electrode 68, producing a current. An appropriate potentialis applied to the electrode 68 from the power source contained inelectronics housing 54 resulting in a signal, which is sent to thesignal processing unit. The signal is then passed through thetransmitter which transmits the information regarding the analyte ofinterest to an external monitor and/or implantable pump as has beendiscussed. The power source, the signal processing unit, and thetransmitter are completely encapsulated in a housing 55 which isimpermeable to biological fluids. The same housing 55 or a separatehousing 70 also encapsulate the analyte sensor except for the membrane62.

The sensor(s) to be incorporated into the device may be eitherelectrochemical, piezoelectric, thermoelectric, acoustic, or optical. Asknown to those skilled in the art, there is a significant body ofliterature regarding the development of electrochemical glucose sensors.These generally incorporate an enzyme, which selectively reacts withglucose.

Electrochemical biosensors may be categorized as amperometric,conductometric, or potentiometric. Amperometric measurements are basedon the oxidation or reduction or electrochemically active substancesinvolved in the oxidation of glucose via glucose oxidase. Another methodis measurement of changes in local pH due to the gluconic acid producedusing a potentiometric sensor, usually a coated-wire pH selectiveelectrode and/or ion-selective field effect transistor (ISFET).Conductometric sensors are based on the principle of electricalresistance changes during the reaction.

Potentiometric and conductometric sensors are currently limited due tothe presence of numerous interfering chemicals in the environment. Themain disadvantage of these sensors is their low sensitivity. Theresponse of the potentiometric sensor depends on logarithmic changes inanalyte concentration.

Microelectronics using ion selective field effect transistors (ISFET's)have been used for measurements of different analytes in body fluids(Erickson, K. A., et al., “Evaluation of a Novel Point-of-care System,the I-Stat Portable Clinical Analyzer”, Clin. Chem. 39(2):283–287 (1993)which is herein incorporated). This allows miniaturization andintegration of the transducer with associated electronic circuitry intoa single chip. However, corrosion of the semiconductor material surfacein saline and in physiological fluids is presently a problem for in vivouse. This problem may be corrected by surface coating or passivation ofthe ISFET. These types of sensors also belong to the class ofpotentiometric sensors, as described above.

Amperometric sensors respond linearly to the analyte concentration. Ifthe limiting processes in signal generation are the enzymatic reactions,the dependence of the signal on glucose concentration is non-linearaccording to Michaelis-Menton kinetics. When the sensor operates in aglucose diffusion-limited mode, the signal is linearly proportional tothe analyte concentration. Amperometric sensors are further subdividedinto three classes:

-   -   1) Based on the production of hydrogen peroxide or consumption        of oxygen.    -   2) Low molecular weight compounds used as mediators of the        electron transfer process.    -   3) Direct electron transfer between the enzyme and the        electrode.

Oxygen-electrode based sensors:

An operational enzyme electrode was first reported by Updike and Hicks(Updike, J. W., and Hicks, J. P., “The Enzyme Electrode,” Nature, 214:986–8, (1967)) based on a description by Clark and Lyons (Clark L. C.,and Lyons, C., “Electrode Systems for Continuous Monitoring inCardiovascular,” Ann. NY Acad. Sci., 102:29–45 (1962)). In this process,the oxygen consumed in the oxidation of glucose is measured. The Clarkoxygen electrode employs a platinum cathode held at a potential ofapproximately (−)0.6 V versus the saturated calomel electrode (S.C.E.),a sufficiently negative potential to reduce oxygen as follows:O₂+4H⁺+4e ⁻→2 H₂OBecause the species being measured in this reaction is a gas,interference from other species in the biological fluid is negligible.However, the signal is determined from reduction of the initial current,making glucose determinations at low concentrations difficult. Further,the system requires the use of a second electrode without any glucoseoxidase, to determine the local oxygen tension for the glucosemeasurement. In addition, it is necessary to insure that there is excessoxygen in the catalytic layer so that the reaction rate is limited bythe glucose. The ratio of blood glucose to oxygen can be as high as 10to 1 in arterial blood, and 100 to 1 in venous blood. (Jaffari 1995) Anoxygen electrode sensor has been described by Armour et al. (1990),which was implanted in the superior vena cava of six dogs for up to 15weeks, with good agreement with standard in vitro assays.

Hydrogen Peroxide based sensors measure hydrogen peroxide productionbased on the oxidation of glucose at potentials above +600 mV vs. SCE.This signal is directly related to the concentration of glucose in thesample.H₂O₂→O₂+2H++2e−

Unfortunately, the high operating potential required can also result inthe oxidation of other chemical species in the blood. This may beovercome by the. use of membranes. Bindra et al. (Bindra, D. S. et al.,“Design and in vitro studies of a needle type glucose sensor forsubcutaneous monitoring,” Anal. Chem., 63: 1692–6 (1991)) reportedglucose detection for up to 10 days in rats with a needle-type sensor,which consisted of GOD immobilized onto cellulose acetate as an innermembrane, and polyurethane as an outer membrane. Moussy et al. (Moussy,F. et al., “Performance of subcutaneously implanted needle-type glucosesensors employing a novel trilayer coating,” Anal. Chem., 65: 2072–7(1993)) used a needle type sensor with a trilayer coating. Nafion wasused as an outer membrane, and poly (o-phenylenediamine) as an innermembrane to reduce interference from small electroactive species. GODwas immobilized between these two layers. As with the oxygen electrodes,the reaction must be limited by glucose, not oxygen.

In amperometric sensors with mediated electron transfer, oxygen as anelectron acceptor is substituted by an artificial mediator, to overcomethe tissue oxygen dependence of amperometric biosensors. Ferrocene andits derivatives are the most commonly used, although hexacyanoferrate(III), tetrathiafuvalene, and ruthenium hexamine (Jaffari, 1995) havealso been investigated. In these sensors, a process involving themediator instead of oxygen takes place:GOD(red)+Mediator(ox)→GOD(ox)+mediator(red)

At the electrode:Mediator(red)→Mediator(ox)The electrochemical oxidation of the reduced mediator occurs at a lowpotential, thus reducing the sensitivity of the sensor to interferingcompounds.

Limitations to the in vivo use of the sensor involve leaching of themediator, and the possible toxicity of the mediator. This has beenaddressed by approaches such as:

-   -   1) Binding of the mediator to high molecular weight compounds,    -   2) Entrapment of the mediator and enzyme in conducting polymer        films,    -   3) Covalent attachment of the mediator to a polymer film,    -   4) Modification of the enzyme with mediator molecules, and    -   5) Use of polymeric mediators.        All of these approaches have been investigated to reduce        mediator leaching, although none have been tested in vivo        (Jaffari, 1995).

An advantage of this type of sensor, because oxygen is not involved inthe signal generation process, is that the sensor signal becomesindependent of the oxygen concentration (Wilkins, 1996).

Amperometric sensors with direct electron transfer are independent ofthe oxygen concentration and involve direct oxidation of glucose by GODat an electrode constructed from conducting organic salts(charge-transfer organic complexes with electron conductivity). Sensorsbased on this principle have been tested in vivo in rats, but little ispresently known about the biocompatibility of such materials.

Examples of sensors that would be suitable for use in this inventioninclude electrochemical sensors described in U.S. Pat. No. 6,001,067,issued to Shults, et al., “Device and method for determining analytelevels”, Dec. 14, 1999, and U.S. Pat. No. 6,212,416, issued to Ward etal, Apr. 3, 2001, “Device for monitoring changes in analyteconcentration”.

Other types of sensors suitable for practicing the present invention,and which also depend on continuous blood glucose transport to thesensor, are fluorescence based sensors described in patents issued toColvin, including U.S. Pat. No. 5,517,313, (issued May 14, 1996), andfurther described in U.S. Pat. No. 5,894,351 issued Apr. 13, 1999, U.S.Pat. No. 5,910,661 issued Jun. 8, 1999, U.S. Pat. No. 5,917,605 issuedJun. 29, 1999, U.S. Pat. No. 6,304,766 issued Oct. 16, 2001, and thefollowing patents issued to Colvin et al., including U.S. Pat. No.6,330,464 issued Dec. 11, 2001, U.S. Pat. No. 6,344,360, issued Feb. 5,2002.

Another type of sensor for which the present invention is well suited,are pressure-based sensors, in which glucose sensitive hydrogels exertpressures which are related to glucose concentrations, described by Hanet al in U.S. Pat. No. 6,514,689, “Hydrogel Biosensor”, issued Feb. 4,2003.

The most commonly used membrane for implantable biosensors ispolyurethane (Jaffari, 1995). Other membranes which have beeninvestigated include cellulose acetate, polypropylene, silicone rubber,and Nafion. These membranes have shown promise in short term monitoring,but long-term monitoring has been more difficult. Davies et al. (Davies,M. L., et al., “Polymer membranes in clinical sensor application. Part1: an overview of membrane function,” Biomaterials, 13: 971–89, (1992))have reviewed extensively the range of polymers used as membranes forbiosensors. Updike reported that immobilization of glucose oxidasewithin a membrane allowed it to accurately measure glucose levels forwell over one year.

In a preferred embodiment, an amperometric electrode is used. Thecharacteristics of such an electrode, such as one available fromMinimed, Inc. (Sylmar, Calif.), is related to the production of hydrogenperoxide in the conversion of glucose to gluconic acid by glucoseoxidase. The noble metal electrode 68 may be connected to any of avariety of RF transceivers. To avoid the use of batteries, which may betoo large for this application, the system may be powered by, and signaltransmission occurs via an inductive link. This requires an inductivecoil (not shown) to be placed both inside the external receiver and aninductive coil (not shown) within the implantable sensor device 10. Anexample of the type of transceiver to be used is employed in the VentakMini IV automatic implantable cardioverter defibrillator (Guidant Corp,Santa Clara, Calif.). The transceiver coil in the preferred embodimentis specifically adapted for use with a stent platform. The size of thecoil, in addition to the number of turns around the implant limits thepower and signal transmission distance between the implant and theexternal receiver/power supply. In order to maximize the diameter of thecoil, the coil is wound around the outside of the implantable sensordevice 10. The transceiver coil is made from an electrically conductivematerial, and is coated to prevent corrosion. The coil may be alignedwith and bonded to the struts of the implantable sensor device 10, inorder to minimize any impact on expansion of the implantable sensordevice 10. Alternatively, the struts of the implantable sensor device 10themselves may serve as the antenna coil. In addition, in order tomaximize signal transmission, the internal and external coils should bealigned so that their major axes are parallel. The external receivercoil should contain a ferrite core to optimize power and signaltransmission. The external receiver and power supply should be designedso that it can be worn on the patient's body, and can be oriented tomaximize the signal from the implanted sensor. This can be done bycustom-placed foam padding to orient the external receiver and powersupply.

Now referring to FIGS. 6A–6C, alternatively, the sensor device may beheld in place within the vessel by any of a variety of anchoring systemsother than a graft or stent. For example, any of a variety of anchoringstructures may be provided with hooks or barbs to engage the vesselwall. Alternatively, semicircular cages or struts may be used, whichonly extend through a partial circumference of the vessel. FIG. 6A showsstruts 74 that can be positioned securely against a blood vessel. Anaxially extending strut 76 connects the circumferential struts 74, andalso could support a sensor. The support 72 can be manufactured in avariety of ways, such as by injection molding a suitable polymericmaterial or by laser cutting from stainless steel, Nitinol or other tubestock. FIG. 6B shows alternate struts 74 that can be secured against thevessel wall. Again, connecting struts 74 is one or more connectingstruts 76. FIG. 6C shows a modification of the device shown in FIG. 6A,where the support 72 is provided with a sheath 180. In the case of Wardet al. (U.S. Pat. No. 6,212,416), they suggest the possible use of thesensor in an artery or vein, but provide no mechanism to secure thesensor within the vessel. If used as described, without such anchoringmechanism, the device would migrate within the vessel, causingsignificant injury to the patient, such as stroke, myocardialinfarction, or pulmonary embolism, all of which are life threatening.

In order to implant an implantable sensor device within the vasculatureof the patient, a catheter-based delivery system is used. Allimplantable sensor devices are formed to include a central lumen with adiameter in the reduced profile sufficient to allow passage of a guidewire and a catheter tip through it. The implantable sensor device ismounted onto a catheter tip, and then collapsed to as small a diameteras possible. The implantable sensor device may be deployed by removal ofa deployment sheath, or other methods, which may be preferred for thespecific type of stent platform being employed. These methods are wellknown in the art. After the implantable sensor device is deployed, thecatheter and guidewire are removed from the (now enlarged) central lumenin the implantable sensor device. The implantable sensor device isplaced such that neither the catheter nor the guidewire adverselyaffects the electronic circuitry.

Preferably, the implantable sensor device is implanted in a relativelylarge artery or vein (>5 mm) in order to minimize the risk that theimplant may occlude the artery. In addition, a healthy artery or veinshould be chosen, so that the device can open completely, and so thatthe flow patterns are normal.

Clinically, it is now accepted practice to place the stent in a parentvessel so that the stent struts cross the ostium of a side branchvessel. This is called “stent jail.” (Pan, M., et al., “Simple andComplex Stent Strategies for Bifurcated Coronary Arterial StenosisInvolving the Side Branch Origin,” Am. J. Cardiol., 83: 1320–25 (1999)).In addition, for stent-grafts for aortic aneurysm repair, investigationis being carried out regarding stent struts which cross the renal arteryostia. An example of such a suprarenal device is the Talent AorticStent-Graft (Medtronic, Inc.). This suggests that it is possible to havea wire (or transducer) which is placed directly across an artery withoutthrombus formation or thrombo-embolization. Thus, in an alternativeembodiment, the sensor or sensing element (i.e., the transducer) can beplaced directly across the path of the flowing blood, on a surface withlow cross-sectional area, such as a wire with a diameter of 0.003″ to0.025″, or a ribbon, oriented with its narrow edge facing upstream. Thetransducer or sensor can be placed on either the proximal, distal, orlateral faces of the wire or ribbon.

As with sensors mounted on or near the wall of the vessel, it isimportant that the sensor be placed across a large vessel with highblood velocity. This will not result in significant thrombus deposition,and any emboli which may result will be of sufficiently small size thatthey will be readily lysed by the patient.

As with stents, there are a multitude of possible designs for a sensingelement which is placed directly across the bloodstream, as will beappreciated by those skilled in the art. The transducer (or sensingelement) could be placed along the surface of a single, straight wire.FIG. 9A shows a cross-sectional view through a stent 14, which includesa sensing element 120 consisting of a wire-like noble electrode 68,which in turn is covered by a gel-enzyme layer 64 and finally by ananalyte-permeable membrane 62. The outside of the analyte-permeablemembrane 62 is bound to a wire-like structure made of a shape memorymaterial 110 such as nitinol, which is either part of or is bonded tothe stent 14. The shape memory material allows the electrode 68 andsensing element 120 to be positioned directly across the lumen of thevessel, as described above. The electrode 68 is connected to the signalprocessing unit (not shown) through a conductor (not shown). Theremainder of the implantable sensor 10, including the power source,signal processing unit, transmitter, and stent, are positioned to beflush against the vessel wall.

FIG. 9B shows an implantable sensor device 10 with a sensing element 130across the cross-section of the stent 14, with sensor circuitry 22. FIG.9C shows an enlarged cross-sectional view of the sensing element 130 ofFIG. 9B, taken along the 9C—9C line. The sensing element 130 consists ofa wire-like noble electrode 68, which in turn is covered by a gel-enzymelayer 64 and finally by an analyte-permeable membrane 62. The outside ofthe analyte-permeable membrane 62 is bound to a wire-like structure madeof a shape memory material 110 such as nitinol, which is bonded to thestent 14 on both ends. The sensing element 130 is connected to thesensing circuit 24 through a conductor (not shown).

FIG. 9D illustrates a transducer 140, consisting of a wire-like nobleelectrode 68, which in turn is covered by a gel-enzyme layer 64 andfinally by an analyte-permeable membrane 62, which is bound to awire-like structure made of a shape memory material 110 such as nitinol,which is bonded to the stent 14. The sensing element 140 is connected tothe sensing circuit 24 through a conductor (not shown).

FIG. 9E shows another configuration with quadruple sensing elements 150,each at right angles to each other, attached to stent 14, with sensorcircuitry 22. The sensing elements 150 are connected to the sensingcircuit 24 through a conductor (not shown). Alternatively, multipletransducers could be positioned to form a wire mesh, provided the meshsize is sufficiently large to permit blood flow without significantthrombo-embolization. In other acceptable configurations, a very smallsensor surface area impedes blood flow, compared to the blood vessel'scross-sectional area.

Referring to FIG. 7A, there is disclosed a deployment catheter 80 whichmay be utilized to deploy a self-expandable stent type sensor support inaccordance with the present invention. The catheter 80 comprises aproximal end 82, a distal end 84 and an elongate flexible tubular body86 extending therebetween. The length of the tubular body 86 will varydepending upon the intended access point and deployment site for thestent sensor. For example, lengths in the area of about 120 cm to about140 cm are typical for use in a coronary artery implantation by way of afemoral artery percutaneous puncture site. Other lengths for differentaccess sites and deployment sites will be apparent to those of skill inthe art in view of the disclosure herein.

The tubular body 86 may be manufactured in accordance with any of avariety of known techniques, such as by extrusion of appropriatebiocompatible polymeric materials. Known materials which are commonlyused for this application include high density polyethylene,polytetrofluroethylene, nylons, and a variety of others known in theart. Alternatively, at least a portion or all of the lengths of thetubular body 86 may comprise a spring coil, solid wall hypodermic needletubing, or braided reinforced wall, depending upon the functionalrequirements of the catheter.

For most applications, the tubular body 86 will be provided with anapproximately circular cross-sectional configuration having an externaldiameter within the range of from about 0.025 inches to about 0.065inches. In accordance with one embodiment of the invention, the tubularbody 86 comprises a multilumen extrusion having an external diameter ofabout 0.042 inches (3.2 f) throughout substantially all of its length.Alternatively, the tubular body 86 can have diameters as large as 12 Fror higher. For percutaneous placement into larger vessels such as theiliac artery. Additional dimensions, materials and manufacturingtechniques are well known in the angioplasty catheter art.

The proximal end 82 is provided with a manifold 88, having a variety ofaccess ports depending upon the desired functionality of the catheter80. In the illustrated embodiment, the manifold 88 is provided with aguidewire port 90 and a deployment wire port 94. Manifold 88 may bemanufactured by injection molding, or other techniques known in the art.

The distal end 84 of deployment catheter 80 is provided with a collapsedsupport structure 96 having a sensor housing 20 thereon in accordancewith the present invention. The support structure 96 is illustrated inits collapsed, low profile configuration, such as for transluminaladvancement towards a placement site. The tubular body 86 may beprovided with an annular recess 98 near the distal end 84, for receivingthe support structure 96. In addition, the tubular body 86 may beprovided with a recess for receiving the sensor housing 20, therebyreducing the collapsed profile of the loaded catheter 80.

The support structure 96 may be constrained in its reduced crossingprofile configuration in any of a variety of ways as has been discussed.In this illustrated embodiment, the support structure 96 is restrainedin its collapsed configuration by a deployment wire 94. Deployment wire94 extends throughout the length of the tubular body 86 through adeployment wire lumen 108, such that a proximal end of the deploymentwire 94 may be proximally retracted by the clinician. The distal end ofthe deployment wire 100 exits the tubular body 86 at a deployment wireport 100, and loops the support structure 96 in one or more loops orslip knots 102 to restrain the support structure 96 in its collapsedconfiguration. Loops or slip knots 102 are configured such that proximalretraction on deployment wire 94 causes the loops or slip knots 102 tobecome untied or otherwise disengaged, thereby releasing the supportstructure 96 so that it expands radially outwardly from its low profileintroduction configuration to its radially enlarged implantedconfiguration. For applications in which the deployment site is removedfrom the percutaneous access site, the catheter 80 is preferablyintroduced over a guidewire as is known in the art. For this purpose, adistal guidewire opening 104 is in communication with the proximalguidewire port 90 by a guidewire lumen 106 extending therebetween.

An example of a similar delivery system is shown in U.S. Pat. No.5,873,906 to Lau, et al. issued Feb. 23, 1999, which is hereinincorporated by reference.

Referring to FIG. 7B, the self-expanding implantable sensor device 96can be deployed from a tubular restraining sheath 108 by pushing a rod110 optionally attached to a disk 112, until the implantable sensordevice 96 is pushed clear of the restraining sheath 108. An example ofsuch a technique is described in U.S. Pat. No. 5,411,551 to Winston etal. (issued May 2, 1995).

Referring to FIG. 7C, another deployment method for a self-expandingimplantable sensor device 96 is shown. The implantable sensor device 96is restrained onto the shaft of the catheter 114 by a sheath 116 and atether line 118. The sheath 116 unfolds when the tether line 118 ispulled, allowing the implantable sensor device 96 to deploy.

Now referring to FIG. 7D, another deployment method for a self-expandingimplantable sensor device 96 is shown. The implantable sensor device 96is restrained onto the shaft of the catheter 120 by two or three or morerestraining prongs 122,124. The restraining prongs 122,124 are retractedwhen one or more deployment wires 126 are pulled, allowing theimplantable sensor device 96 to expand. An example of this can be shownin U.S. Pat. No. 6,024,763 to Lenker et al. (issued Feb. 15, 2000),which is herein incorporated by reference. However, in this patent, therails are only designed to minimize frictional forces between adeployment sheath and the device. They are not actually used as thedeployment mechanism. In the present case, prongs can be used tominimize the delivery profile of the device. Because the sensor andelectronic circuitry may not collapse to a profile as small or ascircular as a stent or stent-graft, it may be more appropriate toposition the electronic components on one side of the catheter, andlayer the collapsed stent on the opposite side of the catheter. Thiscould be achieved by use of delivery prongs, as described.

Referring to FIG. 7E, the implantable sensor device 96 may be deployedby use of a suitable balloon catheter 128 if a balloon expandable stentplatform is used. The balloon 130 is inflated at elevated pressures of 2to 20 atmospheres, and after the implantable sensor device 96 is fullyexpanded, the balloon 130 is deflated and then the balloon catheter 128is withdrawn. Use of such balloon catheters is well known in the art.

Referring to FIG. 7F, the self-expanding implantable sensor device 96can be deployed from a tubular restraining sheath 108 by proximallyretracting the sheath 108 with respect to a push rod 110 optionallyattached to a disk 112, until the implantable sensor device 96 isexposed clear of the restraining sheath 108. The constrained stentsupport structures 14 at either end of the sensor 96, may be held inplace by a release mechanism, 214, which may be a simple knob or hook.Prior to deployment, the atraumatic catheter tip 216 is in directapposition to tubular restraining sheath 108. As the sensor device ispushed clear of the restraining sheath, a small shaft 222 connects thepush rod 110 or disk 112 to the catheter tip 216. A guidewire 218 passesthrough guidewire lumen 220, which may be located eccentrically withrespect to the axis of the catheter. An example of a similar techniqueis described in U.S. Pat. No. 5,411,551 to Winston et al. (issued May 2,1995) which is incorporated herein by reference. However, in Winston,the guidewire passes directly through the center of the catheter,whereas in the present disclosure the guidewire passes eccentricallythrough the delivery catheter. Allowing the guidewire to passeccentrically through the catheter allows the sensor to be hermeticallysealed, and permits a less complex sensor geometry. Additionally, theeccentric guidewire lumen provides a smaller delivery profile for thesensor and its delivery system.

Now referring generally to FIGS. 5A–C, as the blood passes over thesensor 20, glucose diffuses through the semi-permeable membrane 62. Thehousing design should keep the blood shear rate at the apex of thehousing sufficiently high that there is minimal thrombosis on thissurface and any formed thrombus will be of minimal thickness. Thus,diffusion of glucose through the semi-permeable membrane 62 should bethe rate-limiting step, and surface of the device should not becomecovered in a fibrous capsule, as in the case of subcutaneous sensors.The glucose reacts with the glucose oxidase to produce hydrogenperoxide, which further reacts to produce an amperometric signal. Theamperometric signal is converted by an appropriately designed electroniccircuit so that it may be transmitted using the RF transceiver.

In general, the surface of the sensor 20 should be placed on the luminalside of the stent. With the exception of the glucose permeable membrane,which is part of the sensor, the electronic components of the sensorshould be encapsulated in a conformal coating such as a fluoropolymer ora polyamide, and injection-molded or dip-coated in a silicon rubber,polyurethane, or other biocompatible housing material.

The sensor housing should be given a streamlined shape, with graduallysloped transitions at both its proximal and distal ends, in order tominimize flow disturbances. The housing should be as wide or wider atits base than at its apex. The apex of the housing should not protrudehigher than 5–50% of the diameter of the fully deployed device. Theglucose permeable membrane 62 should be placed at the apex of thehousing, should not be encapsulated in the housing material, and shouldface the bloodstream. The glucose permeable membrane 62 should occupythe majority of the area of the housing apex.

Alternatively, if the device includes a graft on the luminal surface ofthe stent, the sensor may be placed on either the luminal or abluminalsurface of a graft material. If the graft is place on the abluminalsurface of the stent, the sensor should be placed on the luminal surfaceof the graft. The graft material is preferably a fluoropolymer, such asePTFE. If the housing is placed on the luminal surface of the graft, itis bonded by partially dissolving the housing material thermally orusing an appropriate solvent so that the housing becomes physicallyinterpenetrated with the graft or with an appropriate adhesive, such asmelt-processed poly(tetrafluoroethylene-co-hexafluoropropylene) (FEP).If the sensor is placed between the stent structure and the graft, itshould be bonded to both the stent and the graft; to prevent graftmovements from causing thrombosis.

Generally, the height of the sensor is estimated to be 2 to 3 times thethickness of a stent strut, based on some observations of a limitednumber of stent designs (Virmani, 1999). This estimate represents theapproximate thickness of the fibrous tissue layer above the stentsurface. Thus, the sensor height above the stent luminal surface shouldbe 1–2 times the thickness of a stent strut. The strut thickness varieswith design, but is roughly 0.005″ to 0.010″ (0.13 to 0.26 mm). Thus, byusing this approach, the housing should protrude about 0.2 to 0.5 mmabove the stent luminal surface.

The effect that this will have on the target vessel is a function of thediameter of the vessel, as given in the following table:

Vessel Diameter Sensor Height above Loss in Loss in Area (mm) struts(mm) Diameter (%) (%) 10 0.5 5.0 9.8 10 0.2 2.0 4.0 9 0.5 5.6 10.8 9 0.22.2 4.4 8 0.5 6.3 12.1 8 0.2 2.5 4.9 7 0.5 7.1 13.8 7 0.2 2.9 5.6 6 0.58.3 16.0 6 0.2 3.3 6.6 5 0.5 10.0 19.0 5 0.2 4.0 7.8As is evident from the table, the smaller the vessel and the greater thesensor height, the greater the obstruction which is created. Designrequirements will vary and require investigation for each application asis well known in the art.

This table also provides insights into the intrusion of the sensor as apercentage of post-deployment vessel diameter. A native vessel withgreater than 50% diameter stenosis is clinically defined to be arestenotic vessel, and typically requires re-intervention. It isestimated that the sensor should not obstruct more than about 25% of thevessel diameter. Alternatively, the percent area loss of a vessel, alongwith a sensor height, can be examined and examples are tabulated below:

Vessel Diameter Sensor Height above Loss in Diameter Loss in Area (mm)struts (mm) (%) (%) 10 0.5 5.0 9.8% 10 2.5 25.0 43.8% 9 0.5 5.0 9.8% 92.3 25.0 43.8% 8 0.4 5.0 9.8% 8 2.0 25.0 43.8% 7 0.4 5.0 9.8% 7 1.8 25.043.8% 6 0.3 5.0 9.8% 6 1.5 25.0 43.8% 5 0.3 5.0 9.8% 5 1.3 25.0 43.8%Using this approach, the sensor height can range from about 0.3 to 2.5mm in height above the stent luminal surface.

With regard to flow velocity, flow measurements in piping systems arecommonly obtained from the pressure drop across a restriction, such as aVenturi meter. As is discussed above, the restriction should not begreater than about 25% of the vessel diameter. Because the volumetricflowrate, Q, must be the same both proximal to and within therestriction, the velocity within the restriction is related to theupstream velocity by:v ₂ =v ₁(A ₁ /A ₂)

Therefore, for an approximately 25% restriction in diameter, therestriction in area is about 43.8%, and the velocity within therestriction is about 178% of the upstream velocity, and for anapproximately 5% restriction in diameter, the velocity within therestriction is approximately 111% of the upstream velocity.

The above analysis is true for situations where the velocity profile isin turbulent flow in a pipe. For laminar flow, there will be a parabolicvelocity distribution, with zero velocity at the wall, and maximumvelocity occurring in the middle. This distribution may be somewhatimpractical to measure, so the exact position where 178% of the proximalvelocity occurs may be hard to establish. Thus, another approach todetermining the proper height of the sensor is to find the sensor heightwhere the blood velocity is approximately 125–200% of its proximalvelocity. This can be determined using duplex ultrasound, or hot-wireanemometry, or other flow-measuring techniques. Using this approach, thesensor may be flush mounted against the wall of the vessel, with a flowimpedance device mounted at the same axial position within the vessel,in order to increase the velocity within the sensor/impedance device toabout 125–200% of the proximal flow velocity. The location where theestimate of the proximal flow velocity is approximately 200% ispreferred. This would allow for the possibility of flush mountedsensors.

One key problem with most implanted sensors is fibrous tissueencapsulation. While proper positioning of the sensor within the vesselcan minimize the thickness of a fibrous tissue layer, it may not bepossible to avoid endothelialization of the sensor surface. Thus, thesensor design of the current invention may not actually projectcompletely beyond the tissue growth, but the thickness of the tissuelayer would ideally be only a single layer of endothelial cells.

Prior to implantation, the sensor may be checked for a reproducibleresponse to glucose concentration. This may be done in the operatingtheater using sterile technique immediately prior to implantation, ormay be done in a batch-wise manner during the manufacturing process.Following implantation, the sensor may then be calibrated by comparisonof the output signal with an accepted standard. Typically, this may bedone by comparing the signal with the value obtained from a laboratoryglucose sensor, such as made by Yellow Springs Instruments (YellowSprings, Ohio). Preferably, the calibration curve should be stable overtime and for wide range of glucose values. For example, the slope of thecalibration curve should be sufficiently stable to give an error of lessthan ten percent. Weekly calibrations should be sufficient to insurestable and accurate readings, however calibration can be performed asfrequently as required. For example, there may be a change insensitivity over the first month or so, if the transducer becomesendothelialized. However, after that point, the system should be stable.Thus, calibrations could be required as often as every other day atfirst, and taper off to no more than about once per week.

Once calibrated, if the external signal produced by the sensor indicatesthat the glucose level is outside of the normal physiological range,there are several possible outcomes.

-   -   1. An audible, visible or tactile alarm may sound, so that the        patient or physician may check the sensor with a home glucose        monitoring kit (e.g., One Touch, LifeScan, Johnson & Johnson),        and then take appropriate action such as administration of        insulin or glucose.    -   2. The signal may be transmitted directly to an implantable        insulin pump, which may administer insulin directly without        requiring a response by the patient.

FIGS. 5A–C show various embodiments in which the sensor and transmitterare on either the luminal or abluminal surface of the stent. Nowreferring to FIG. 5A, an implanted sensor 20 is shown in a transversecross-sectional view through the vessel. The struts 60 of theimplantable sensor device may be surrounded by an inner tubular sheath(not illustrated for simplicity), which would contact the blood vesselwall when deployed. The sensor housing 20 sits between a pair of struts60. The membrane 62 is exposed to the blood flow. The analyte sensor 56will normally have a larger cross-sectional area than the stent struts60. The outer sheath allows for enhanced uniform radial expansion(beyond that of self-expanding struts), especially if the sheath isbonded to each of the struts 60 of the stent, and is bonded through thelength of the stent, not just at the ends. This would link each of thestruts 60 to its neighbors, preventing uneven expansion. It would alsobe advantageous to make the sheath out of a material which could bestretched slightly to obtain its final diameter, so that irregularitiesin the flow surface are minimized, except in the region of the sensor.

When a non-sheath implantable sensor device is either balloon expanded(or in the case of self-expanding stents, following balloon touch-up),the stent struts 60 can be embedded more deeply into the vessel wallthan the sensor housing 20. If the struts 60 were positioned between theflowing blood and the sensor surface, they would cause flow stagnation,and therefore thrombosis on the membrane 62 of the sensor. If the sensoris placed instead on the luminal surface of the stent, the sensor willagain be embedded less deeply in the vessel wall, although withoutstruts 60 on its luminal surface, there will be minimal hemostasis andthrombus formation on the transducer surface.

As it can be seen, the sensor could be placed either between the stentstruts and the inner sheath, or on the luminal surface of the innersheath. In both cases, a semi-permeable membrane might still benecessary to insure that only the analyte of interest reaches thesurface of the sensor. In either case, the sensor should be designedwith a streamlined profile at both its proximal and distal ends, tominimize regions of hemostasis.

In accordance with another aspect of the invention, the sensor and itsassociated circuitry are connected between two or more stent segments,with or without the presence of a stent or supporting member at thelocation of the sensor. This provides a number of advantages, includinga decrease in the delivery profile, and an increase in flexibility. Thisallows easier access to sites with tortuous vascular anatomy, but willstill allow the sensor to maintain its same relative position in theblood vessel.

FIG. 10 shows a side elevational view of a sensor 20, a sensing circuit24 and signal processing unit 26, a power source 28 and aradio-transmitter 30, with expanded anchoring stents 14 at the proximalend 200 of the sensor 20, intermediate between the sensor 20 and thesensing circuit 24 and signal processing circuits 26, also intermediatebetween the sensing circuit 24 and signal processing circuits 26 and thepower source 28 and radio-transmitter 30, and at the distal end 210 ofthe power source 28 and radio-transmitter 30. Each segment of theimplantable sensor device is connected to its neighboring segment by theuse of electrical connectors 57, with strain relief elements 226. Theelectrical connectors 57 are hermetically sealed. Additionally, theelectrical connectors 57 are of sufficient length, and have redundanciesso that they can connect the various segments allowing flexibility ofpositioning between each segment of the implantable sensor whileminimizing the strain imparted on the connectors. It is important tonote that the relative position of each of these components between theanchoring stents is somewhat arbitrary. The device could be designedwith these components in any order, and still function equally well, aswill be apparent to those skilled in the art.

FIG. 11 shows a side elevational view of a sensor 20, a sensing circuit24 and signal processing unit 26, a power source 28 and aradio-transmitter 30, with anchoring stents 14 in the compressed statesuch as within a tubular deployment catheter at the proximal end 200 ofthe sensor 20, intermediate between the sensor 20 and the sensingcircuit 24 and signal processing circuits 26, also intermediate betweenthe sensing circuit 24 and signal processing circuits 26 and the powersource 28 and radio-transmitter 30, and at the distal end 210 of thepower source 28 and radio-transmitter 30. The advantages of thissegmented embodiment of the invention are that the overall deliveryprofile of the device is reduced, and the device becomes more flexiblefor delivery through tortuous vessels.

Thus, the present invention provides at least one electrical component,having at least a first support on a proximal end and a second supporton a distal end, for supporting the component in a body lumen. Asdisclosed herein, the electrical circuitry may desirably be divided intoat least two or three discrete physical components spaced axially apartand in electrical communication with each other to minimize theimplanted profile and enhance deliverability. Each of the components maybe provided with a proximal and a distal support, as illustrated inFIGS. 10 and 11. The two or three or four or more supports may beconnected together independently of the intervening electricalcomponent, or may be connected only through the housing of theintervening electrical component. The supports may comprise any of avariety of self-expandable coils or other structures well known in theself expandable stent and graft (e.g., abdominal aortic aneurysm graft)arts. Alternatively, balloon expandable or mechanically expandableanchor structures may be used. In an alternate configuration, thesupport may comprise an elongated ribbon or wire such as Nitinol whichis based into a spiral, having the electrical components spaced axiallyapart therealong.

In another aspect of this invention, it is possible to monitorsubstances other than glucose by using different enzymes, as listed inthe table below:

ANALYTE ENZYME glucose glucose oxidase glucose glucose dehydrogenaselactate lactate oxidase 1-methionine 1-amino acid oxidase1-phenylalanine 1-amino acid oxidase d-aspartate d-amino acid oxidased-glutamate d-amino acid oxidase urate urate oxidase ethyl alcoholalcohol oxidase methyl alcohol alcohol oxidase cholesterol cholesteroloxidase ascorbic acid ascorbate oxidase

In still another aspect of the invention, a different class of sensorsis used to detect the presence of chemical analytes in blood. Thesesensors, termed “immunosensors”, rely on the interaction between anantibody and its antigen, which are very specific, and typically thereis a very strong interaction between antibody and antigen. This type ofdetection system may have broader applicability for detecting moleculesin blood as compared with enzymatic sensors (Rabbany S Y, Donner B L,Ligler F S, “Optical Immunosensors” Crit Rev Biomed Eng 1994;22(5–6):307–46; and Stefan R I, van Staden J F, Aboul-Enein H Y,“Immunosensors In Clinical Analysis” Fresenius J Anal Chem March–April2000; 366(6–7):659–68). Antibodies can be produced which can recognizealmost any biomolecule, and therefore the number of possible targetanalytes could be substantially increased using this method. As Stefanet al. (2000) state, “The main problem of [immunosensor] utilization isthe interference or loss of affinity when real biological fluids (e.g.,blood, serum, plasma, urine, saliva) have to be analyzed.” Theycontinue, “For in vivo tests with immunosensors, highly biocompatiblematerials have to be found for electrode construction” (emphasis added).The present invention provides an excellent method for solving the keyissue of immunosensor fouling.

There are a number of potential applications of this technology,including detection of infectious disease, cardiac disorders, andcancer, as well as clinical drug monitoring. (Rabbany 1994). Otherpotential applications include illicit drug monitoring, and researchapplications focused on drug development and pharmacokinetics studies.For instance, Hanbury et al. (Hanbury C M, Miller W G, Harris R B,“Antibody Characteristics For A Continuous Response Fiber OpticImmunosensor For Theophylline” Biosens Bioelectron 1996; 11(11):1129–38)describe a continuous immunosensor for monitoring theophylline, avasodilator with a narrow therapeutic range (55–110 μM), and whichrequires frequent monitoring to assure therapeutic efficacy and preventtoxicity.

Another important application would be in the area of crisis-orienteddiagnostics. For example, heart patients presenting with chest pain, orthose at risk for recurrent acute myocardial infarction (AMI), could bequickly diagnosed with AMI by using the present invention to monitor forbiochemical markers such as creatine kinase (CK-MB), serum cardiactroponins (cTnT or cTnI), aspartate aminotransferase (AST), lactatedehydrogenase (LDH), beta-hydroxybutyrate dehydrogenase (HBD), serummyoglobin, glycogen phosphorylase isoenzyme BB (GPBB), fatty acidbinding protein (FABP), phosphoglyceric acid mutase isoenzyme MB,enolase isoenzyme alpha beta, S100a0, and annexin V (Olukoga A,Donaldson D, “An Overview Of Biochemical Markers In Acute CoronarySyndromes”, J Royal Soc Promot Health 121(2):102–106 (2001).) Ofpotentially greater value is use the present invention to monitor forbiochemical markers which may precede AMI, such as C-reactive protein(CRP) (Rabbany 1994), serum TnT, and inflammatory markers such as V-CAM,I-CAM, or interleukin-6 for the prognosis of myocardial infarction. Forsuch applications, speed and accuracy of monitoring are essential.

Other potential uses of this type of immunosensor include monitoringanticoagulation levels, which might include monitoring for thrombin,prothrombin fragment 1+2 (F1+2), fibrinopeptides A or B, Factor Xa,thrombin-antithrombin III complex, or platelet release products such asthromboxane A2, PDGF, or markers of fibrin degradation such as D-dimer,other markers which could potentially serve as an index for a patient'slevel of anticoagulation. Finally, the prospect of continuous,ambulatory monitoring provides improved therapeutic efficacy, costcontainment, convenience, and patient peace of mind.

Continuous immunosensors have been previously demonstrated (Ligler, etal., U.S. Pat. No. 5,183,740 “Flow Immunosensor Method And Apparatus,”Feb. 2, 1993; and Ligler, et al., U.S. Pat. No. 6,245,296 “FlowImmunosensor Apparatus,” issued Jun. 12, 2001) for the detection ofexplosives in soil samples, among other applications (Gauger P R, Holt DB, Patterson C H Jr, Charles P T, Shriver-Lake L, Kusterbeck A W,“Explosives Detection In Soil Using A Field-Portable Continuous FlowImmunosensor,” J Hazard Mater May 7, 2000; 83(1–2):51–63; Vianello F,Signor L, Pizzariello A, Di Paolo M L, Scarpa M, Hock B, Giersch T, RigoA., “Continuous Flow Immunosensor For Atrazine Detection,” BiosensBioelectron Jan. 1, 1998; 13(1):45–53; Narang U, Gauger P R, KusterbeckA W, Ligler F S, “Multianalyte Detection Using A Capillary-Based FlowImmunosensor,” Anal Biochem Jan. 1, 1998; 255(1):13–19; Kusterbeck A W,Wemhoff G A, Charles P T, Yeager D A, Bredehorst R, Vogel C W, Ligler FS, “A Continuous Flow Immunoassay For Rapid And Sensitive Detection OfSmall Molecules,” J Immunol Methods Dec. 31, 1990; 135(1–2):191–7; andCharles P T, Conrad D W, Jacobs M S, Bart J C, Kusterbeck A W,“Synthesis Of A Fluorescent Analog Of Polychlorinated Biphenyls For UseIn A Continuous Flow lnnunosensor Assay,” Bioconjug ChemNovember–December 1995; 6(6):691–4.) and sample analysis is very rapid(<5 minutes). However, the continuous immunosensors described are notreadily translated to the clinical need for ambulatory monitoring. Forinstance, Ligler (1993) teaches that the level of the target molecule inthe system should be measured by determining the amount of label that isreleased from the apparatus, not the amount of antibody-antigen complexremaining in region of the apparatus. Measuring the amount of thelabeled target molecule released into the bloodstream would require thatsamples be taken from the patient on a frequent basis. This would notprovide any benefit over using a standard enzyme-linked immunosorbentassay (ELISA) or radioimmunoassay (RIA) to look for this same unlabeledanalyte of interest.

As described in FIGS. 1–5 and 9–11, above, the sensor 20 may in apreferred alternative embodiment, be an immunosensor. The anchoringplatform for this embodiment of the invention is similar to that whichwas previously described in the context of electrochemical glucosesensors, as is the positioning of the sensor 20 and radiotransmitter 30with respect to the anchoring platform 14, and the radiotransmitter 30for sending the signal to the external device. A power source 28,sensing circuit 24, and a signal processing circuit 26 are also includedwith the immunosensor. The external device provides a read-out of thedata regarding the concentration of analyte in the patient's blood atany time, and appropriate audible, visual, vibratory, or other signalsare given to notify the user of an important change in condition.

As with enzymatic glucose sensors, there are a multiplicity of ways inwhich the molecular recognition event (i.e., the antibody-antigenreaction) can be translated into an electrical signal. These includeamperometric, potentiometric, piezoelectric, surface plasmon resonance(SPR), scintillation, acoustic, fluorescent and chemiluminescentimmunosensors (Stefan 2000). These types of sensors have been previouslydescribed in the literature (Stefan 2000).

Amperometric sensors typically use enzymes (Rabbany 1994) such asalkaline phosphatase or horse-radish peroxidase to label antigens orantibodies, so that the reaction products produced by the enzyme (ratherthan the antibody or antigen) provide the actual signal to be detected(Stefan 2000). This adds a level of complexity for an implantabledevice, because a third component is now required, in addition to theantibody and antigen. For this type of immunosensor, not only are theantigen of interest and an enzyme-labeled antibody required, but anappropriate substrate for the enzyme label is also required. This enzymesubstrate can be delivered under conditions to insure that the signalbeing produced is determined by the amount of enzyme-labeled antibodypresent, rather than the amount of substrate for the enzyme.

Potentiometric sensors have excellent reproducibility, althoughaccording to Stefan (2000), “in most cases, potentiometric transducerscannot provide the necessary sensitivity for the antigen-antibodyreaction.” This is because the resulting signal is proportional to thelogarithm of the analyte concentration.

Piezoelectric sensors, such as those based on the quartz-crystalmicrobalance (QCM) may not be well suited to this application, due tothe high background response from non-specific adsorption (Stefan 2000).Surface plasmon resonance detection is also less well suited to thepresent application, due to interference from non-specific binding(Rabbany 1994, Stefan 2000).

In principal, detection could be performed by the use of antigenslabeled with radionuclides. In that case, a scintillation detector suchas a NaI crystal coupled with a photodiode or photomultiplier tube couldbe used. The detector for this type of sensor is otherwise similar tothat described below for the fluorescence detectors. However, the use ofradionuclides as labeling agents is less attractive, due to problemswith handling, and exposure of the patient and physician to radiation.

Optical immunosensors include fluorescent or chemiluminescence-basedsensors. A fluorescence-based sensor is shown in FIG. 12, and issupported by a stent structure 14 which is in the compressed state fordelivery from a catheter 120. In this sensor, the labeled targetmolecule 260 has a fluorescent label, such as fluorescein,tetramethylrhodamine, or Texas Red. If a fluorescent label is used, thesensor will consist of both a light source 280 such as a light-emittingdiode (LED), and a photosensitive detector such as a photomultipliertube or photodiode 290, such as that supplied by Silicon Sensors(Madison, Wis.). The light source 280 with a filter (not shown) produceslight at the excitation wavelength λ₁, and the photodiode 290 with afilter (not shown) detects the light at wavelength λ₂ emitted by thefluorescence of the label. In the case of fluorescein, the wavelength oflight that is absorbed by the label is in the range of 470 nm, and thewavelength of fluorescent light is in the range of 540 nm. The sensingcircuit 24 and radiotransmitter 30 are also included. The intensity oflight detected by the photodiode provides an electrical signal, which isdependent upon the intensity of the detected fluorescence. A change insignal level will indicate a change in the concentration or presence ofthe antigen of interest.

FIG. 12 also shows a sensor 20 containing a reservoir 300 of labeledantigen 320 which is included as part of the sensor. The reservoir isencapsulated on all but one side by a membrane 330 that is impermeableto the analyte of interest. The impermeable membrane may be apoly(carbonate urethane), silicone rubber, Parylene, Teflon, orwater-impermeable polymer. A semi-permeable membrane 62 is included onthe final side of the reservoir, which adjoins the sensor compartment asa semi-permeable barrier to control the rate at which the labeledantigen diffuses from the reservoir to the sensing compartment. Thesemi-permeable membrane 62 is preferably a hydrogel, such aspoly(ethylene glycol) (PEG), poly(vinyl alcohol) (PVA), poly(N-vinylpyrrolidone) (PVP), polyacrylamide, poly(acrylic acid) (PAA),poly(hydroxyethyl methacrylate) (PHEMA), or other. The thickness,chemical nature, and cross-link density of the hydrogel can becontrolled as is known by those skilled in the art in order to obtain anappropriate rate of diffusion of labeled antigen out of the reservoir,and into the sensor compartment. Further, this semi-permeable membrane62 may contain multiple layers, such as a hydrogel bonded to a dialysismembrane, polyurethane, poly(vinylidene fluoride) (PVDF), ePTFE, Nafion,or other second membrane layer. The reservoir provides a flux of labeledtarget molecule which allows the sensor to be replenished over time, andallows the measurement of both decreases and increases in analyteconcentration in the blood. The reservoir provides the added benefit ofextending the lifetime of the sensor.

The sensor for this embodiment detects the presence of the targetmolecule as follows. First, an antibody 260, which hereafter refers toany of combination of polyclonal antibodies, monoclonal antibodies, orthe Fab fragment of an antibody, is immobilized on or near the sensorlight source 280 and photodetector 290. The antibodies may be fromhuman, mammalian, or non-mammalian origins, provided that adequatecross-reactivity between the antibody and antigen can be established.The antibodies may be immobilized within or upon films (not shown) ormembranes (not shown) that may be present on the surface of the lightsource 280 or photodetector 290. In addition, the antibodies may beimmobilized onto particulate supports near the sensor component, such asSepharose (Pharmacia). The immobilization may be performed by covalentlybonding the antibody to the substrate with bi-functional molecules suchas glutaraldehyde, carbodiimides, biotin-avidin, and other moleculeswith one or more functional groups on each of at least two ends as arewell known to those skilled in the art. Additionally, bi-functionalspacer molecules such as N-hydroxysuccinimide derivatized polyethyleneglycols may be used to bind the antibody within the sensor compartment.Because immobilization of antibodies can significantly change theirreactivity, methods for improved orientation of the antibodies have beendescribed (Stefan, 2000). During the manufacture or final preparation ofthe device, prior to implantation in the patient, the antibody issaturated with a large excess of labeled antigen, beyond thestoichiometric excess needed to fully occupy all the available sites onthe antibodies.

In comparison with enzymatic electrochemical sensors, immunosensors maypreferably be modified so that the luminal sensor surface is positionedeven closer to the center of the blood vessel in order to further reduceendothelial coverage, because not all antigens are capable of passingthrough an endothelial layer. The outer membrane 62 covering the luminalsurface of the sensor is preferentially a hydrogel, such aspoly(ethylene glycol) (PEG), poly(vinyl alcohol) (PVA), poly(N-vinylpyrrolidone) (PVP), poly(hydroxy ethylmethacrylate) (PHEMA), or other,in order to further reduce the probability of cell adhesion to thesurface. It may also come from the group consisting of ePTFE,polyurethane, silicone rubber, cross-linked collagen, polypropylene,cellulose acetate, poly(vinylidene fluoride) (PVDF), Nafion or otherbiocompatible material. The outer membrane must also be permeable to theanalyte of interest. The permeability of the membrane is selected toallow the analyte of interest to freely contact the sensor, whilerestricting the passage of other blood components. The permeability ofthis membrane 62 can be controlled by varying the porosity of thehydrogel or polymer, which can be controlled by varying the cross-linkdensity and the molecular weight of the polymer between cross-links.Additionally, the thickness of the outer membrane can be controlled,allowing control of the transport rate of the target analyte. Finally,the outermost membrane may be bonded to a second, inner membrane layerwhich may be made from either ePTFE, polyurethane, silicone rubber,cross-linked collagen, polypropylene, cellulose acetate, poly(vinylidenefluoride) (PVDF), Nafion, or other biocompatible membrane material withcontrolled porosity to further control the transport of the targetmolecule to the sensor.

As the sensor is exposed to flowing blood containing the antigen ofinterest, there will be a competition for the antibody binding sites onthe sensor between the labeled antigen (both that originally implantedin the sensor compartment 270 and that contained in an additionalreservoir 300 of labeled target molecules) and the antigen in thebloodstream. Over time, there will be a displacement of the labeledantigen originally supplied in the sensor by the non-labeled antigen,which is present in the bloodstream. Rabbany (1994) reports that theantibody-antigen reaction is seen as nearly irreversible except in thecase of continuous-flow immunosensors. Because of this, only whenunlabeled antigen is present in the continuously flowing blood is thelabeled antigen displaced. Thus, if no antigen is present in the blood,the sensor will remain undepleted. The amount of antigen-antibodycomplex 310 can be measured on a very frequent (nearly continuous)basis. To do this, the LED emits a pulse of light, and the resultingfluorescence intensity is measured by the photodiode.

Now referring to FIG. 13, the light emitting diode sensor circuitconsists of a constant-current LED drive circuit 404 that ensuresconsistent brightness over a range of input voltages. The LED 402 iscontrolled by a square wave clock circuit 406 that gates the LED 402 onand off at rate that can be set for the particular application. Thistype of sensor is powered by the power supply 408 via the powercoil/antenna 410.

The light detector circuit consists of a photodiode 412 and acurrent-to-frequency converter 414. The I-F converter output is gated bythe clock 180 degrees out of phase with respect to the signal used todrive the LED. This ensures that there is no output from the converterwhen the LED is being driven.

In an alternate embodiment, the reservoir can instead be fabricated froma degradable material such as poly(lactide-co-glycolide) (PLGA),polycaprolactone (PCL), or other controllably degradable material. Thelabeled target molecule is incorporated into such a matrix by mixing thetarget molecule directly into a solution of PLGA, and precipitating thePLGA. Alternatively, the reservoir is loaded with microspherescontaining either the antibody or antigen of interest (Ma J, Luo D, QiW, Cao L, “Antitumor Effect Of The Idiotypic Cascade Induced By AnAntibody Encapsulated In Poly(D,L-Lactide-Co-Glycolide) Microspheres,”Jpn J Cancer Res 92(10); 1110–15, (2001); Torche A M, Le Corre P, AlbinaE, Jestin A, Le Verge R, “PLGA Microspheres Phagocytosis By Pig AlveolarMacrophages: Influence Of Poly(Vinyl Alcohol) Concentration, Nature OfLoaded-Protein And Copolymer Nature,” J Drug Target 7(5):343–54 (2000);Mordenti J, Thomsen K, Licko V, Berleau L, Kahn J W, Cuthbertson R A,Duenas E T, Ryan A M, Schofield C, Berger T W, Meng Y G, Cleland J,“Intraocular Pharmacokinetics And Safety Of A Humanized MonoclonalAntibody In Rabbits After Intravitreal Administration Of A Solution Or APLGA Microsphere Formulation,” Toxicol Sci 52(1):101–6, (1999)). Thesemicrospheres may be prepared by emulsion polymerization, as is known inthe art. It is possible to control the release rate of labeled targetmolecule from these degradable microspheres by varying the co-polymercomposition and particle size. For example, a 50:50 ratio of lactide toglycolide monomers is known to degrade more rapidly than homopolymers ofeither poly(lactide) or poly(glycolide). One difficulty with these typesof systems is that as the matrix degrades, its by-products can affectthe pH of the local micro-environment. This may have an effect on thestability and solubility of the labeled antigen. A solution to this maybe found by incorporating basic salts, such as magnesium hydroxide,calcium hydroxide, calcium phosphate, or zinc sulfate, in order tomaintain a pH neutral environment, as described by Zhu et al (Zhu G, andSchwendeman S P, “Stabilization of Proteins Encapsulated in CylindricalPoly(lactide-co-glycolide) Implants: Mechanism of Stabilization by BasicAdditives”, Pharm Res 17(3):351–7 (2000)). Another approach by Jiang(Jiang W, and Schwendeman S P, “Stabilization and Controlled Release ofBovine Serum Albumin Encapsulated in Poly(D,L-lactide) and Poly(ethyleneglycol) Microsphere Blends”, Pharm Res 18(6): 878–885 (2001)) is toblend a highly water-soluble polymer such as poly(ethylene glycol)(PEG), poly(N-vinyl pyrrolidone) (PVP), or poly(vinyl alcohol) (PVA),with a more slowly degrading homopolymer, such as poly(lactide) (PLA) orpoly(glycolide) (PGA). However, the only requirement for the presentinvention is that the microsphere delivery system is able to maintainantigenic recognition of the labeled target molecule. It is notnecessary that the activity of the target molecule be maintained.

One of the difficulties with measuring the amount of labeledantigen-antibody complex still present near the sensor, however, is thatof sensor response time. That is, if the amount of label bound to thesensor is relatively large compared to the amount which is displaced, itwill be difficult to accurately determine the concentration of theanalyte. This is because the difference between two large and uncertainnumbers also has a high degree of uncertainty. Thus, the currentapproach will have a somewhat longer sensor response time, or require alarger change from baseline, and hence have a somewhat lowersensitivity, as compared to the approach described by Ligler (1993), inorder to establish a measurable change in antigen level. Sensor responsetimes are dependent on the concentration of antigen in the bloodstream,the rate at which the antigen is transported across the semi-permeablemembrane, the reaction rate between antigen and antibody, and the rateat which the displaced, labeled antigen is transported out of the sensorcompartment, into the bloodstream. In addition, the relativeconcentrations of labeled antigen from the reservoir and unlabeledantigen from the bloodstream, which are present in the sensorcompartment, will also play an important role. Similarly, the lifetimeof the sensor will be dependent on the factors listed above, as well asthe amount of labeled target molecule incorporated in the reservoir. Ifthe sensor lifetime were limited by the amount of labeled targetmolecule incorporated in the reservoir, it may be possible to extend thesensor lifetime, for example, by re-loading the sensor withcontrolled-release microspheres in a secondary catheterizationprocedure, or by direct injection with a syringe and needle.

In an alternative embodiment, it is possible to immobilize the antigen,and label the target antibody, in order to monitor the presence ofantibodies in the bloodstream. This is done by reversing theimmobilization of the antigen near the sensor, and providing a reservoirof labeled antibody as described above for FIG. 12. This may be usefulfor monitoring patients with conditions such as HIV or AIDS.

The implantable sensor described, could also be made to be retrievable.(Neuerburg, J, Gunther, R W, Rassmussen, E, Vorwerk D, Tonn K, Handt S,Kupper W, Hansen J V, “New Retrievable Percutaneous Vena Cava Filter:Experimental In Vitro And In Vivo Evaluation,” Cardiovasc InterventRadiol 16(4): 224–9 (1993); Neuerburg, J M, Gunther, R W, Vorwerk D,Dondelinger R F, Jager H, Lackner K J, Schild H H, Plant G R, Joffre FG, Schneider P A, Janssen J H, “Results Of A Multicenter Study Of TheRetrievable Tulip Vena Cava Filter: Early Clinical Experience,”Cardiovasc Intervent Radiol 16(4): 224–9 (1993); and Millward S F, OlivaV L, Bell S D, Valenti D A, Rasuli P, Asch M, Hadziomerovic A, Kachura JR, “Gunther Tulip Retrievable Vena Cava Filter: Results From TheRegistry Of The Canadian Interventional Radiology Association,” J VascInterv Radiol 12(9): 1053–8 (2001).) At any point up to a biologicallydetermined retrieval limit, it is possible to remove both the sensor andits anchoring platform, using a device such as the GooseNeck Snare(Microvena Corp.) in a follow-up catheterization procedure.

As shown in FIG. 14A, the anchoring platform 14 can be designed with ahook, loop, eye, or other easily snareable feature 250 such as at thedownstream end, so that it can be removed up to a certain period of timewithout the necessity of leaving the device permanently implanted in thepatient. In this example, the hook is preferably made of the samematerial as the anchoring platform, such as 316L stainless steel, ornitinol, and may either be laser-cut from the same tubing stock as theanchoring platform, or may be laser-welded onto the anchoring platform.The hook may preferably have a diameter greater than that of theretrieval snare wire, typically 0.020–0.030″ in diameter, so that theretrieval snare can easily rest in the hook. As an alternative to ahook, a feature such as a ball with a diameter of about 0.75 to 2 mm maybe formed at the end of straight segment of wire with a diameter ofabout 0.5 mm. The hook may also be shaped like a closed loop, or like ananchor, with each arm of the anchor having a length of about 0.5 to 1mm. In FIG. 14A, the snareable feature 250 on the anchoring platformprotrudes slightly into the lumen of the vessel, so that it does notbecome completely encapsulated in tissue. In an alternative embodiment,as shown in FIG. 14B, the snareable feature 250 is mounted to the distalend of the sensor housing 20. Typically in order to prevent corrosion,the sensor electronics will be hermetically sealed using either a glass,ceramic, epoxy, or metal housing. In this case, the snareable featuremay be either laser-welded to the metal housing, or embedded in theglass, ceramic or epoxy at the time the hermetic seal is formed.

The sensor housing may be coupled to the anchoring platform usingmechanical clips designed to release under the appropriate appliedforce, or using degradable materials such as PLGA or PLA. Because theluminal surface of the sensor housing is designed so that it does notbecome encapsulated with fibrous tissue, it is possible to retrieve thesensor component at any time following implantation. However, withdegradable materials, the implant must remain in place long enough forthese degradable anchors to lose their strength and give way duringretrieval. The anchoring platform is left behind. In these embodimentsof the invention, the snareable feature makes the device suitable forcontinuous monitoring of a temporary patient condition. It is alsosuitable for short-term monitoring of drugs, such as those with a narrowwindow between efficacy and toxicity. In addition, it is possible toremove a sensor that may no longer be functional.

Now referring to FIGS. 15A and 15B, in an alternative embodiment, boththe sensor and its anchoring platform could be designed to beretrievable for an indefinite period. In this case, the sensor 20 iscentrally mounted in an anchoring platform 400. The sensor housing maybe bonded to the platform by positioning the two components in closeproximity during the formation of a hermetic seal, using glass, metal,or epoxy, as described above for FIG. 14A. Additionally, if the sensoris hermetically sealed in a metal case, the sensor and anchoringplatform could be laser-welded together. Another means of connecting thetwo components would be by mechanical interfit. The anchoring platform400 contains side struts 410 which extend axially within the vessel. Theanchoring platform 400 is preferably made of a self-expanding materialsuch as nitinol. This anchoring platform 400 centers the sensor 20 inthe middle of the flowstream, so that the proximal (upstream) tip of thesensor is again kept free of thrombus or other fouling tissue because ofthe high flow velocity. The side struts 410 also help to keep the sensorcentered in the bloodstream. These side struts 410 may have a lengthranging from 1 to 100 mm, with longer struts providing improved abilityto prevent the sensor from tipping away from the center of the vessel,and shorter struts making the entire device easier to introduce into thevasculature, and also being easier to retrieve. A preferred length rangefor the side struts 410 is from 5 to 30 mm. The side struts 410 maycontain radially outwardly directed hooks (not shown) to prevent thedevice from migrating. These migration-resistant hooks could be placedat any position along the side struts, and are oriented in a directionto prevent migration of the device towards the heart or lungs. Asdescribed in U.S. Pat. No. 6,258,026, to Ravenscroft et al., issued Jul.10, 2001, the hooks could be made small enough so that they preventsensor migration, but can be easily deformed during retrieval, allowingthe device to be withdrawn from the vessel. The proximal tip of thesensor is preferably in a streamlined configuration, such as a parabolaor a cone, to minimize the risk of thrombus formation. A hook 250, knob,or other easily snareable feature is placed at the distal (downstream)end of the sensor. The entire device can then be removed in acatheterization procedure by using a snare as previously described tocatch the hook 250 and draw the anchoring platform into the distal endof a tubular retrieval catheter. Because the parallel struts 410 of theanchoring platform 400 are open-ended, they will not become mechanicallyinterlocked by neointimal tissue, and therefore the entire device willbe retrievable for an indefinite period. The retrieval process would bemuch like withdrawing a hypodermic needle from under the skin.

It is possible to remove both the sensor and its anchoring platform,using a snare in a follow-up catheterization procedure. As is well knownto those skilled in the art, the snare would be introduced into thevasculature using the same Seldinger techniques that are used to implanta stent. Under fluoroscopic guidance, the physician would first selectan appropriate vascular access site, and then would insert a guidingcatheter of sufficient diameter so as to be able to accommodate theretrieved sensor and its anchoring platform. Next, the physician wouldinsert a snare, such as the GooseNeck Snare (Microvena Corp.) throughthe guiding catheter, and approach the sensor hook with the snare. Oncethe physician grasps the sensor hook with the snare, and proximallyretracts the snare, the sensor would collapse into the retrievalcatheter, and the sensor and guiding catheter could be simultaneouslywithdrawn from the patient's body.

There are at least two types of sensors which are used in monitoringarterial blood gas values, and which are suitable for practicing thisinvention. The first type is a Clark-type electrode, in which anelectrochemical reaction, such as the reduction of O₂, occurs at anelectrode, and an electrical current is monitored. Early oxygenelectrodes required initial calibration against a simultaneous ABGsample, and sensor drift necessitated frequent recalibrations.Thrombosis was also a critical issue for this type of intravascularsensor, and so the oxygen electrode was not widely accepted. (Mahutte,1998). In addition, electrochemical pH and pCO₂ monitors are notavailable.

The other major category of IABG's is optodes, which measure blood gasesoptically, and are suitable for practicing the present invention. Thesesensors are based on fiberoptics that are coupled with a reagent.Interaction between the reagent and an analyte of interest results in achange in the optical properties of the reagent, which is detectedthrough the fiber-optic. Optical sensors such as these do not requirethe use of a reference electrode. There are three main types of optodes,based on either absorbance, fluorescence, or quenching (W. R. Seitz,“Chemical Sensors Based on Fiber Optics”, Anal. Chem. 56(1):16A–34A(1984) and C. K. Mahutte, “On-line Arterial Blood Gas Analysis withOptodes: Current Status,” Clin. Biochem 1998; 31:119–130). When light ofa certain wavelength passes through a reagent, and there is a reductionin intensity of the light, that is termed absorbance. Fluorescenceoptodes work because the reagent absorbs light at one frequency, andre-emits light at a second frequency. The intensity of light at bothfrequencies can be measured, providing a correlation with theconcentration of the analyte of interest (L. A. Saari and W. R. Seitz,“pH sensor based on immobilized fluoresceinamine”, Anal. Chem. 1982;54:821–23). Fluorescence quenching is also used for detection ofanalytes such as oxygen. Oxygen can reduce the amount of fluorescence ofcertain organic compounds, and therefore the reduction in the intensityof re-emitted light can be used as the basis for measurement of oxygenconcentration (J. I. Peterson et al., “Fiber-Optic Probe for In VivoMeasurement of Oxygen Partial Pressure”, Anal. Chem. (1984) 56:62–67)).

The reagent phase of an optode can be positioned at any point along thefiber-optic, either at the end, or along the sides, allowing fordifferent geometric configurations to be used (Mahutte 1998).

There have been a number of ABG sensors that have reachedcommercialization at various times and these would be suitable forpracticing the present invention. These include the Paratrend 7(Pfizer/Biomedical Sensors, Malvern, Pa.) (I K Weiss, et al, “Continuousarterial gas monitoring: Initial experience with the Paratrend 7 inchildren,” Intensive Care Med. (1996) 22:1414–1417), the CDI 1000(Cardiovascular Devices, Irvine Calif.), Optex Biomedical (TX), and thePB3300 (Puritan-Bennett, Carlsbad, Calif.) (T. Lumsden, et al. “ThePB3300 Intraarterial Blood Gas monitoring system,” J. Clin. Monit 1994;10:59–66).

The Paratrend 7 is a combination electrode-optode system, withfiberoptic pH and pCO₂ sensors, an amperometric oxygen sensor, and athermocouple for temperature compensation, and would be suitable forpracticing the present invention. The CDI 1000, the Optex sensor, andthe PB3300 are pure optode systems. In the PB3300, the reagent is on theexternal circumference of the sensor. Acceptable levels of accuracy havebeen reported with the PB3300 system, but the longest reported durationof implantation was 121 hours.

In addition, U.S. Pat. No. 5,326,531 Jul. 5, 1994 to Hahn et al., “CO₂Sensor using a hydrophilic polyurethane matrix and process formanufacturing” describes a CO₂ optode, and U.S. Pat. No. 5,378,432 Jan.3, 1995 to Bankert et al, “Optical Fiber pH microsensor and method ofmanufacture” describes a pH optode, both of which can be used in thepresent invention.

The clinical issues identified which prevent the sensors fromfunctioning properly in vivo are known as the “wall-stress effect”, andare attributed to hypotension or vasoconstriction. In some instances,the pH, pCO₂ and pO₂ values are variable due to thrombus formation atthe tip of the catheter, while in other instances, the variation was dueto the sensor touching the arterial wall, and measuring gas values inthe tissue (C. K. Mahutte, “On-line Arterial Blood Gas Analysis withOptodes: Current Status,” Clin. Biochem 1998; 31:119–130). Bothvasospasm and wrist flexion have been identified as factors whichnegatively affect the performance of these sensors. The anchoringplatform (14) shown in FIG. 16 is designed to address these problems.

In U.S. Pat. No. 6,447,395 (Nov. 5, 2002), Schulman et al. shows asensor designed to avoid direct contact between the sensing surface andthe vessel wall. This sensor is built near the tip of a catheter, andthe catheter at the sensing surface is bent in a zig-zag shape. However,there are two major drawbacks with this type of design. First, bloodflow stagnation will occur in the region at the inside apex of thezig-zag, and therefore, thrombosis will occur on this sensing surface,and interfere with sensor function, even if the sensor surface does nottouch the vessel wall. Second, because most catheters are made fromsomewhat rigid materials, such as nylon, any catheter with a zig-zagpattern as shown in FIG. 6 of Schulman would not be able to beintroduced into a vessel using a catheter sheath introducer, as isstandard clinical practice. In addition, such a sensor could not beintroduced into a vessel using a trocar or other small diameter tool toallow vessel access.

As shown in FIG. 16, the anchoring platform 14 is connected to acatheter 500, which contains multiple sensors 20, such as optodesensors, for monitoring pH, pO₂, and pCO₂, near its distal end. As inother embodiments, the anchors are made of a self-expanding materialsuch as nitinol. As previously described, these anchors help to positionthe sensor near the wall, with the sensing surface 20 oriented towardthe vessel lumen. These anchors may be split rings, which under acompressive load would form closed rings, resisting vessel compressiondue to vasospasm (Khatri, S. et al., “Stenting for Coronary ArterySpasm,” Catheter Cardiovasc Interv., May 2000, 56(1):16–20, and Cheng,T. O., “Percutaneous Coronary Intervention for Variant Angina: Balloonvs. Stent”, Catheter Cardiovasc Interv., 2002 May 56(1):21) but wouldalso open up under tension to allow the catheter and sensor to bewithdrawn at the end of its useful life. For monitoring of blood gasesin the radial artery, which has a typical diameter of 3–4 mm, theseanchors should have an expanded diameter of 3.5–5 mm, to allow forslight oversizing of the anchors, which is common practice in stentplacement. The catheter and sensor may be withdrawn from the patientusing a slightly larger catheter sheath, in order to accommodate theanchors, and protect the vessel from further damage during removal. Asan alternative to placing the sensor near the side of the vessel, thecatheter and sensor could be centered in the vessel, with one or morestruts connecting the sensor to one or more support rings. The supportrings serve to maintain the position of the sensor in the center of thevessel, and additionally serve to resist vessel compression due tovasospasm, and may reduce the risk of vessel compression due to jointflexion. In another embodiment, the support rings may have a helicalconfiguration, the helix being joined to the sensor by one or morestruts, which help to position the sensor in the center of thebloodstream. The use of a helical support structure which is connectedto the catheter at only a single point provides the option of removingthe sensor, even after the wound healing process has occurred, andfibrin which deposits on the strut surfaces has transformed intofibro-collagenous tissue. This is because the helix can be unwound, andpulled into the retrieval catheter. Thus, the addition of these supportanchors serves to counteract vessel compression, one of the mainobstacles that prevents IABG sensors from functioning long-term in anarterial setting. Appropriate positioning of the sensor in the flowfield also minimizes the risk that slow flow will cause thrombosis andsensor fouling. Alternatively, the sensors may be placed in an arterythat is larger than the radial artery, such as the femoral artery, andthe present invention is suitable for numerous intravascular locations.

In addition to monitoring pH, O2, and CO2, it is also possible to useoptodes to measure any of a variety of different species usingion-selective membranes with ionophores. (Bakker, E., et al.,“Carrier-Based Ion-Selective Electrodes and Bulk Optodes. 1. GeneralCharacteristics” Chem. Rev. 1997 97:3083–3132). These ion selectivesensors, whether optical or electrochemical, are based on lipophiliccomplexing agents capable of reversibly binding ions, called ionophores.The sensing layer of most ion selective electrodes (ISE's) or optodes isan organic polymeric membrane, of which polyvinyl chloride (PVC) is themost widely used. Other materials such as derivatized PVC, siliconerubber, or polyurethanes can also be used. A requirement for suchmembranes is that its glass transition temperature needs to be below theoperating temperature. This requirement is necessary to allow reasonablyrapid diffusion times. Alternatively, plasticizers, such as those usedin PVC can be used to achieve the same goal. The plasticizer istypically selected for its compatibility with the ionophore. Theionophore determines the selectivity of the sensor. The final componentof the ISE membrane is either a quaternary ammonium salt, to providecationic sites, or tetraphenylborate salts to provide anionic sites. (R.D. Johnson et al, “Ionophore-based ion-selective potentiometric andoptical sensors,” Anal. Bioanal. Chem. 2003 June; 376(3):328–41.)

It is possible to alter the surface of these ISE membranes in order toimprove their blood compatibility, such as with the use of phosphorylcholine (Berrocal M J, et al, “Improving the blood-compatibility ofion-selective electrodes by employing poly(MPC-co-BMA), a copolymercontaining phosphorylcholine, as a membrane coating.” Anal. Chem. Aug. 12002; 74(15):3644–8) or with heparin, poly(ethylene glycol), orpoly(vinyl pyrrolidone), or the ISE membrane may first be covered by aprotective membrane such as ePTFE with 0.4 micron porosity, and thensurface modifications applied. Among the species that can be monitoredusing these technologies are calcium, chlorine, potassium, sodium,bicarbonate, phosphate, phosphorous, and magnesium. Other metal ionswhich can be detected using this technology include lithium, ammonium,rubidium, cesium, beryllium, strontium, barium, molybdenum, iron,copper, silver, zinc, cadmium, mercury, thallium, bismuth, lead,uranium, and samarium. Other inorganic ionic analytes which can bemeasured using this technology include creatinine, organic ammoniumions, nucleotides, and polyionic analytes such as heparin and protamine.(Buhlmann P, et al, “Carrier-Based Ion-Selective Electrodes and BulkOptodes. 2. Ionophores for Potentiometric and Optical Sensors,” ChemRev. Jun. 18 1998; 98(4):1593–1688). The response of ISE's toward asingle analyte ion is described by the Nernst equation.

In a further aspect of the invention, a pressure sensor can be implantedwithin a blood vessel in order to monitor a patient on a continuousbasis. The pressure sensor may be similar to many described in the art,such as potentiometric pressure sensors, inductive pressure sensors,capacitive pressure sensors, piezoelectric pressure sensors, and straingage pressure sensors. The pressure sensor can also be a microelectromechanical system (MEMS) device such as those made by Honeywell,Sensym, or All Sensors Corp (San Jose, Calif.) that is constructed usingmicrofabrication methods that are well known in the microelectronicsindustry. Once the pressure signal has been converted into an electricalsignal, the signal is then transmitted to an external receiver that maybe worn by the patient. The sensor can be configured on the anchoringplatform in any of the various configurations previously described inthis disclosure, including those in which the sensor itself isretrievable in the case of sensor malfunction.

It is also valuable to monitor the condition of patients who haveabdominal aortic aneurysms, which can be treated with endovascularstent-grafts. Sonesson et al. (Sonesson B., et al., “Intra-aneurysmpressure measurements in successfully excluded abdominal aortic aneurysmafter endovascular repair” J Vasc Surg 37:733–8 (2003)) describe howmonitoring the pressure on both the luminal side and aneurysmal side ofsuch stent-grafts can help a physician to determine when and whetherfollow-up intervention is required. However, in their study, a follow-upcatheterization procedure was required to monitor these pressures. Itwould clearly be advantageous to be able to obtain this informationwithout having to perform an additional procedure. An implantablepressure sensor as described herein which does not become covered with alayer of fibrocollagenous tissue would provide accurate intra-arterialmeasurements. If a second sensor were mounted on the aneurysmal side ofthe stent-graft, its signal would not change substantially, as thehealing process within the aneurysmal sack is known to be substantiallydelayed.

Schmitz-Rode et al. (Schmitz-Rode T, Schnakenberg U, Pfeffer J G, PirothW, Vom Bogel G, Mokwa W, Gunther R W, “Vascular capsule for telemetricmonitoring of blood pressure,” Rofo Fortschr Geb Rontgenstr NeuenBildgeb Verfahr. 2003 February; 175(2):282–6) describe a capsularpressure sensor that is implanted in the arterial system at an arterialbranch. Three self-expanding legs were attached to the sensor capsule inorder to maintain sensor position and prevent migration as well asocclusion (WO 00/74557, 7 Jun. 2000 to T. Schmitz-Rode, et al.).However, they reported that in 3/6 cases, the devices migrated severalcentimeters to the next arterial bifurcation, where the device becamelodged, and caused an occlusion of one of the branch vessels. Thethree-leg anchoring system of this device means that it is prone totilting, and so its exact position with the vessel is not controllable.Further, this device is not easily retrievable in the event of patientneed or device failure. In addition, physicians would prefer to use adevice whose placement is more accurate. It may also be important tomonitor pressures at a location distant from an arterial bifurcation.Hence, an improved anchoring system, as described herein, can provide asignificant improvement in this area.

Bullister et al. (Bullister E., et al., “A Blood Pressure Sensor forLong-Term Implantation,” Artificial Organs 25(5):376–379 (2001))describe a pressure sensor that is mounted within the wall of a titaniumtube, and that may be used in line with a left ventricular assistdevice. However, such a device would require surgical implantation,which may not be appropriate for monitoring many patients, especiallycritical care patients, whose condition may exclude them as surgicalcandidates. In addition, if such a sensor were anastamosed end-to-end ina vessel, there would be a clear compliance mis-match between the vesseland the sensor. Compliance mis-match is well established as a cause ofvascular graft failure.

Pressure sensors that are attached to stents have also been disclosed(U.S. Pat. No. 6,015,387 issued Jan. 18, 2000 to Schwartz et al. andU.S. Pat. No. 6,053,873 issued Apr. 25, 2000 to Govari et al.). Govaridescribes positioning of a pressure sensor on the outer wall of a stent.However, without proper positioning of the sensor with respect to thestent and the vessel lumen, the sensor surface will be covered withfibrin, which will subsequently transform into fibrous tissue, as partof the natural healing process. The fibrous tissue will create a barrierbetween the sensor and the blood, and will affect the value of thepressure readings obtained from the sensor. The fibrous tissue itselfhas elastic properties that will vary over time, and will be dependenton the thickness of the tissue layer that is formed. Thus, such a systemmay not allow immediate determination of whether the sensor is coveredwith tissue, or whether patient intervention is required. In addition,the devices described by Schwartz and Govari are not designed to beretrievable in case the sensor fails to function properly.

Ultrasonic flow sensors of the type described in U.S. Pat. No. 6,053,873to Govari et al., which is hereby incorporated in its entirety, may bemounted in a hermetically sealed sensor housing made from epoxy, glass,ceramic, stainless steel, or the like, and mounted on the luminalsurface of a stent, such that the luminal surface of the sensor housingprotrudes into the lumen, as previously described. The sensor mayalternatively be placed between two or more anchors, in such a mannerthat the sensor housing is designed to protrude into the lumen of thevessel, with respect to the anchors. The sensors are mounted atspaced-apart locations so that the transit time or Doppler ultrasoundmethods for determining flow rates described by Govari may be used.Additionally, the sensors may be positioned at an appropriate Dopplerangle.

The anchors may consist of self-expansive materials, such as nitinol, ormay be made from balloon expandable materials such as 316L stainlesssteel. In addition the anchors may be made from elements that are notclosed curves, and which do not become interlocked with fibrous tissueover time. The open hoop structures shown in FIG. 16 are one type ofanchor that may be suitable for practicing this invention. There mayoptionally be provided a retrieval hook at the downstream end of thesensor for removing the sensor at the end of its useful life. The sensoritself may be powered using an inductive link, or an implantablebattery, and the sensor responses may be analyzed using anapplication-specific integrated circuit.

A thermal sensor for measuring flowrates is disclosed which consists ofa thermal source, such as a resistive heating element, and at least onesensing unit, such as a thermistor or thermocouple. Both the heatingelement and the sensor are placed in the sensor housing, which protrudesinto the lumen of the vessel in order to minimize tissue overgrowth. Theheating element is placed upstream from the first thermocouple, and asmall (<2.5° C.) temperature increase is applied to the bloodstream nearthe heating element. A second thermocouple is positioned immediatelydistal to the heating element to insure that the local temperature riseis small. The time for the temperature increase to be sensed by thedownstream thermocouple is measured, and the flowrate may be determinedfrom this information using standard thermodilution calculations.

Additionally, the Fick method for determination of cardiac output couldbe performed, using oxygen sensors described herein, placed in both thearterial circulation, and in the pulmonary artery. The Fick method fordetermining cardiac output is based on the principle that consumption ofa substance (oxygen in this case) must equal blood flow to the organmultiplied by the difference between the arterial and venousconcentrations of the substance. For this method, the formula forcardiac output is as follows:

${CO} = \frac{{oxygen}\mspace{14mu}{consumption}\mspace{14mu}{per}\mspace{14mu}{minute}\mspace{14mu}({VO2})}{( {{{arterial}\mspace{14mu}{oxygen}\mspace{14mu}{content}} - {{venous}\mspace{14mu}{oxygen}\mspace{14mu}{content}}} )}$where oxygen content is calculated as: (1.34×[Hb]×oxygensaturation)/100. [Hb] is the hemoglobin concentration, which can beperiodically assessed by a hemotology laboratory. In this case, theoxygen consumption can either be estimated or directly measured usingstandard techniques, or using the implantable sensors described herein.Arterial oxygen saturation is usually determined by arterial blood gasanalysis (or using an implantable oxygen sensor described herein), whilevenous oxygen saturation is determined by mixed venous (pulmonaryarterial) blood gas analysis (or using an implantable oxygen sensordescribed herein). There are a number of different sensors that would beappropriate for the measurement of nitric oxide (Berkels, R. et al. Anew method to measure nitrate/nitrite with a NO-sensitive electrode. JAppl Physiol 90: 317–320, 2001). Because NO has a short half-life (2–30s), it is difficult to measure authentic NO. It rapidly decomposes tonitrate and nitrite, which may accumulate in the sample. There aredifferent methods of determining NO. It is possible to measure NO byusing a bioassay, an oxyhemoglobin assay, electron paramagneticresonance, chemiluminescence, HPLC, the Griess reaction, or differentelectrochemical electrodes. In addition, electrodes for NO detection arecommercially available (WPI, Sarasota, Fla., and Inter Medical Co.,Ltd., Nagoya, Japan).

Kilinc E, et al (J Pharm Biomed Anal. 2002 Apr. 15; 28(2):345–54)describe the use of an electrochemical sensor, similar to a glucoseelectrode, in that it consists of a amperometric, bi-polymer modified,platinum-iridium microelectrode. Most common biological interferencessuch as ascorbic acid, uric acid and glucose were eliminated viabi-polymer coatings of four layers of Nafion and a layer of 50 mMo-phenylenediamine (OPD).

The electrochemical reaction occurs at a working electrode at an appliedvoltage of 900 mV vs. Ag/AgCl electrode. The reaction is describedbelow:NO-e⁻ NO⁺NO⁺+OH⁻ HNO₂The redox current flowing between the working and reference electrodesis proportional to the concentration of NO, and is measuredamperometrically (Zhang, X, et al., “Amperometric Detection of NitricOxide”, Mod. Asp. Immunobiol. 1(4), 160–65, 2000).

An alternative system for measuring NO, involves the use of a microchipsensor (Zhang, X, et al., “A Novel Microchip Nitric Oxide Sensor withsub-nM Detection Limit”, Electroanalysis 2002; 14(10):697–703. Thissystem is also commercially available and may be employed in the currentinvention.

The sensor may be factory-calibrated using a nitric-oxide donorcompound, such as SNAP (Zhang, Mod Asp Immunobiol 2000). Additionally,if deemed appropriate for the patient, calibration may be performedfollowing implantation by local administration of a NO-releasing agent,such as nitroglycerin. The system may not require a high degree ofaccuracy, because there is a large burst of NO during ischemic stroke,and as long as such a burst was detected, the exact concentration levelmay not be critical.

In a preferred embodiment, an outer coating of heparin is applied to alayer of poly(ethylene glycol) (PEG), poly(vinylpyrrilidone) (PVP), orother hydrogel, in order to minimize thrombosis or fouling. The hydrogelis applied to a bioprotective layer of ePTFE, which prevents attack bymacrophages. A polymer such as poly(dimethylsiloxane) or a polyestermembrane such as Sympatex (Azko Nobel, Wuppertal, Germany) which ispermeable to NO is employed. (Wang C, Deen, W M, “Nitric oxide deliverysystem for cell culture studies”, Ann Biomed Eng. 2003 January;31(1):65–79.) The ePTFE surrounds the NO-permeable membrane, and theNO-permeable membrane surrounds the electrode. A porphyrinic sensor maybe constructed by coating a carbon-fiber electrode with a solution ofnickel (II) tetrakis (3-methoxy-4-hydroxyphenyl) porphyrin (Tschudi, M Ret al, “Direct In Situ Measurement of Nitric Oxide in MesentericResistance Arteries”, Hypertension 1996; 27:32–35), and then applyingthe outer membranes, excluding the NO-permeable membrane. Additionally,a layer of drug-releasing polymer, as described below, may be placedbetween the ePTFE membrane and the NO-permeable membrane, or may beplaced only in certain areas between the ePTFE and NO-permeablemembranes, such that there is direct contact between these twomembranes. The coating or layer chosen to minimize thrombosis or foulingshould have a porosity large enough to allow the analyte of interest toreach the sensing surface. In certain applications, radial inwarddisplacement of the sensing surface may be reduced and longevity of thesensor still exist.

In addition to monitoring NO levels, it is also possible to measurenitrite or nitrate concentrations, using ion-selective electrodes, aspreviously described (Johnson, R D, et al, Anal Bioanal Chem (2003) 376:328–341 and Bakker, E, Chem. Rev. 1997, 97, 3083–3132). Ion selectiveelectrodes for monitoring nitrate and nitrite ions are commerciallyavailable (Vernier Software & Technology 13979 S. W. Millikan Way,Beaverton, Oreg. 97005 or Techne® Inc., 3 Terri Lane, Suite 10,Burlington, N.J. 08016). In addition, it is possible to measure nitriteion concentrations by electrochemical oxidation of nitrite to nitrate,which occurs at a similar potential to the oxidation of nitric oxide tonitrite. Therefore, in another preferred embodiment, measurement ofnitric oxide and nitrite and nitrate ions are performed simultaneously,by the use of nitric oxide and nitrite electrodes and nitrite andnitrate ion-selective electrodes.

Itoh Y, et al (Anal Biochem. Dec. 15 2000; 287(2):203–9), describe afluorescent measurement for nitric oxide (NO) using diaminofluoresceinDAF-FM and its diacetate (DAF-FM T). DAF-FM is converted via anNO-specific mechanism to an intensely fluorescent triazole derivative(DAF-FM T). They showed that in the presence of 1 μM DAF-FM, theconcentrations of NOR-1, an NO donor, in the range of 2–200 nM waslinearly related to the fluorescence intensity. However, this reactionis somewhat slower than the use of an NO electrode, having a maximalresponse at about 30 minutes.

This fluorescence reaction for monitoring NO can be monitored by the useof optodes, as described above for arterial blood gas monitoring, andfor the immunosensor, described herein. An excitation light source at awavelength of 500 nm is passed through the fiberoptic and reacts withthe DAF-FM. The emission wavelength is at 515 nm. A photodiode detectsthe fluorescence intensity, and sends a signal via the radiotransmitterto an external alarm device worn by the patient. Alternatively, thealarm system can be set to contact someone besides the patient viatelephony, such as a relative or physician, to notify them thatimmediate action needs to be taken.

While it is anticipated that hemodynamic effects will significantlyreduce biological fouling, in the case of NO detection, it mayadditionally be desirable to further insure the prevention ofendothelial growth by the elution of certain pharmacologic agents toprevent endothelial coverage of the sensor surface. Such anti-angiogenicagents include: paclitaxel (Belotti D, et al., “Themicrotubule-affecting drug paclitaxel has antiangiogenic activity”, ClinCancer Res. 1996 November; 2(11):1843–9.), rapamycin (Novak, K.,“Swinging the vote for rapamycin” Nature Reviews Cancer 2002 2:75.)docetaxel (Guo, X L, et al., “Inhibitory effects of docetaxel onexpression of VEGF, bFGF and MMPs of LS174T cell” World J Gastroenterol2003; 9 (9): 1995–1998.) TNP-470 (Yeh, J R J, et al, “The antiangiogenicagent TNP-470 requires p53 and p21CIPyWAF for endothelial cell growtharrest”, PNAS, 2000 97(23):12782–12787), carboxyamido-triazole (CAI),thalidomide, or interleukin-12 (Masiero, L, et al. “Newanti-angiogenesis agents: review of clinical experience withcarboxyamido-triazole (CAI), thalidomide, TNP-470, and interleukin-12”Angiogenesis 1997; 1(1):23–35.).

These may be released from a degradable matrix, such as PLGA, or from anon-degradable matrix, such as a ethylene vinyl acetate (EVA). Thisdrug-releasing layer may be incorporated within the sensor housinginside the outer bioprotective layers.

In an alternative embodiment, it may be desirable to incorporate anenzyme in an enzyme-gel layer, as discussed previously with regard toglucose sensors. However, in this case, the enzyme is nitrate reductase,which converts nitrate into nitrite. The nitrite is thenelectrochemically oxidized back to nitrate, and provides an electricalcurrent that is measured amperometrically.

In an alternative embodiment, as shown in FIG. 18, the nitric oxide ornitrite or nitrate sensor 20 may be incorporated into the distal end ofa catheter-like tube 700. The proximal end of the catheter-like tube isattached to a hermetically sealed enclosure 710 which contains theantenna 30 and/or battery 720, and the electronic circuitry. Wire leads730 are then used to connect the sensor 20 with the electroniccircuitry. The device is implanted in a patient in a manner similar totransvenous pacemaker insertion which is well known to those of skill inthe art. The physician makes a small incision in the skin, and thencreates a hole in the vein to insert the distal tip of the cathetercontaining the sensor using techniques known to those of skill in theart. The hole is preferably smaller than the diameter of the distal tipof the catheter so that the distal tip is forced to stay in place. Thedistal tip containing the sensor is then advanced into the vein eitherblind or using ultrasound guidance. The enclosure containing theelectronics is then placed subdermally, and the insertion site is closedusing standard surgical techniques known to those of skill in the art.The signal indicative of the nitric oxide or nitrite or nitrate presentis transmitted to a device worn externally by the patient, as describedpreviously, and can communicate to the patient by an alarm which iseither auditory or vibratory, or in the event of a stroke, can transmitinformation by telephony to either a physician, relative, caregiver, oremergent care provider.

In an alternative embodiment, it is possible to retrieve the sensorwithout the presence of a hook on the sensor or anchoring platform.Preferably, this is done using a sensor housing that is bonded to theanchoring platform using a degradable material, and this procedure isperformed after the bonds have degraded. Also, this procedure ispreferably performed using a distal protection device, such as theAngioguard (Cordis Corp.). Because the sensor is not covered by a layerof neointimal tissue, it is possible to remove the sensor using theretrieval system shown in FIGS. 17A and 17B. A catheter 600 with a setof clips 640 is shown, and a pair of balloons 620, 630 is placed on theside of the catheter tip opposite the clips. The catheter is maneuveredinto position under fluoroscopic guidance, so that the clips 640 areadjacent to the sensor. Next, balloons 620 and 630 are inflated eithersequentially or simultaneously, as appropriate for the individual case,so that the clips are forced around the sensor housing. Next, balloons620 and 630 are deflated. Finally, the sensor that is being removed, andthe catheter tip are separated from the wall of the vessel by inflatingballoon 610. Balloon 610 is then deflated, and the entire system isremoved from the patient, along with the sensor.

Although the present invention has been described in connection withcertain preferred embodiments, persons of ordinary skill in the art willunderstand that many modifications can be made thereto within the scopeof the claims which follow. Accordingly, the scope of the invention isnot intended to be limited by the above description but instead is to bedetermined entirely by reference to the following claims.

1. A sensor for implantation within a blood vessel comprising: a supportstructure; a sensor housing carried by the support structure; and asensing surface exposed to the exterior of the housing; wherein thesensor is configured to detect nitric oxide or a metabolic of nitricoxide, and wherein the sensor has streamlined configuration with respectto the support, and wherein the sensing surface includes a layer thatminimizes the formation of thrombus.
 2. An implantable sensor as inclaim 1, wherein the support structure comprises a stent.
 3. Animplantable sensor as in claim 1, wherein the support structurecomprises a catheter.
 4. An implantable sensor as in claim 1, whereinthe support structure comprises an expandable metal mesh.
 5. Animplantable sensor as in claim 1, wherein the sensor housing ispositioned on the luminal side of the support structure.
 6. Animplantable sensor as in claim 1, wherein the sensor housing ispositioned within an opening on the side wall of the support structure.7. An implantable sensor as in claim 1, further comprising a tubularsleeve surrounding the support structure.
 8. An implantable sensor as inclaim 7, wherein the tubular sleeve is on the radially outwardly facingsurface of the support structure.
 9. An implantable sensor as in claim7, wherein the tubular sleeve comprises ePTFE.
 10. An implantable sensoras in claim 1, wherein the sensor comprises an ion-selective electrode.11. An implantable sensor as in claim 1, wherein the sensor is selectedfrom the group consisting of amperometric electrodes, porphyrinicelectrodes, and microchip electrodes.
 12. An implantable sensor as inclaim 1, further containing an analyte permeable membrane and an enzymegel layer.
 13. An implantable sensor as in claim 12, wherein the enzymegel layer comprises nitrate reductase.
 14. A sensor for implantationwithin a blood vessel comprising: a support structure; a sensor housingcarried by the support structure; and a sensing surface exposed to theexterior of the housing and wherein the sensing surface includes a layerthat minimizes the formation of thrombus; wherein the sensor isconfigured to detect a metabolic product of nitric oxide.
 15. Animplantable sensor as in claim 14, wherein the support structurecomprises a stent.
 16. An implantable sensor as in claim 14, wherein thesupport structure comprises a catheter.
 17. An implantable sensor as inclaim 14, wherein the support structure comprises an expandable metalmesh.
 18. An implantable sensor as in claim 14, wherein the sensorhousing is positioned on the luminal side of the support structure. 19.An implantable sensor as in claim 14, wherein the sensor housing ispositioned within an opening on the side wall of the support structure.20. An implantable sensor as in claim 14, further comprising a tubularsleeve surrounding the support structure.
 21. An implantable sensor asin claim 20, wherein the tubular sleeve is on the radially outwardlyfacing surface of the support structure.
 22. An implantable sensor as inclaim 20, wherein the tubular sleeve comprises ePTFE.
 23. An implantablesensor as in claim 14, wherein the sensor comprises an ion-selectiveelectrode.
 24. An implantable sensor as in claim 14, wherein the sensoris selected from the group consisting of amperometric electrodes,porphyrinic electrodes, and microchip electrodes.
 25. An implantablesensor as in claim 14, further containing an analyte permeable membraneand an enzyme gel layer.
 26. An implantable sensor as in claim 25,wherein the enzyme gel layer comprises nitrate reductase.